Grating-based sensor combining label-free binding detection and fluorescence amplification and readout system for sensor

ABSTRACT

A grating-based sensor is disclosed that has a grating structure constructed and designed for both evanescent resonance (ER) fluorescence detection and label-free detection applications. Some embodiments are disclosed which are optimized for ER detection in an air mode, in which the sample is dry. Other embodiments are optimized for ER detection in liquid mode, in which the sample is suspended in liquid medium such as water. One and two-dimensional gratings are also disclosed, including gratings characterized by unit cells with central posts, central holes, and two-level, two-dimensional gratings. A readout system for such sensors is also disclosed. One embodiment includes a first light source optimized for collecting label-free detection data, a second light source optimized for collecting ER fluorescence amplification data, and at least one detector. In one embodiment, the detector is an imaging system and includes a CCD camera for collecting both ER and label-free data. In other embodiments, the at least one detector takes the form of a spectrometer for collection of label-free data and a photomultiplier for collecting ER data. In other embodiments, a single light source such as a tunable laser or broad band light source is used.

PRIORITY

This application is a divisional of U.S. Ser. No. 11/490,556 filed Jul.20, 2006 now U.S. Pat. No. 7,863,052, which claims priority benefitsunder 35 U.S.C. §119 (e) to the following United States provisionalpatent applications, the entire contents of which are incorporated byreference herein:

(1) Ser. No. 60/707,579 filed Aug. 11, 2005

(2) Ser. No. 60/713,694 filed Sep. 2, 2005

(3) Ser. No. 60/778,160 filed Feb. 28, 2006

(4) Ser. No. 60/790,207 filed Apr. 7, 2006.

BACKGROUND

A. Field of the Invention

This invention relates generally to grating-based biochemical sensordevices and detection instruments for such devices. Grating-basedsensors are typically used for optical detection of the adsorption of abiological material, such as DNA, protein, viruses or cells, smallmolecules, or chemicals, onto a surface of the device or within a volumeof the device. The sensor of this invention has a grating structure thatis constructed in a manner for use in two different applications: (a)label-free binding detection, and (b) fluorescence detection, forexample wherein the sample is bound to a fluorophore or emits nativefluorescence.

B. Description of Related Art

1. Label-Free Detection Sensors

Grating-based sensors represent a new class of optical devices that havebeen enabled by recent advances in semiconductor fabrication tools withthe ability to accurately deposit and etch materials with precision lessthan 100 nm.

Several properties of photonic crystals make them ideal candidates forapplication as grating type optical biosensors. First, thereflectance/transmittance behavior of a photonic crystal can be readilymanipulated by the adsorption of biological material such as proteins,DNA, cells, virus particles, and bacteria on the crystal. Other types ofbiological entities which can be detected include small and smallermolecular weight molecules (i.e., substances of molecular weight<1000Daltons (Da) and between 1000 Da to 10,000 Da), amino acids, nucleicacids, lipids, carbohydrates, nucleic acid polymers, viral particles,viral components and cellular components such as but not limited tovesicles, mitochondria, membranes, structural features, periplasm, orany extracts thereof. These types of materials have demonstrated theability to alter the optical path length of light passing through themby virtue of their finite dielectric permittivity. Second, thereflected/transmitted spectra of photonic crystals can be extremelynarrow, enabling high-resolution determination of shifts in theiroptical properties due to biochemical binding while using simpleillumination and detection apparatus. Third, photonic crystal structurescan be designed to highly localize electromagnetic field propagation, sothat a single photonic crystal surface can be used to support, inparallel, the measurement of a large number of biochemical bindingevents without optical interference between neighboring regions within<3-5 microns. Finally, a wide range of materials and fabrication methodscan be employed to build practical photonic crystal devices with highsurface/volume ratios, and the capability for concentrating theelectromagnetic field intensity in regions in contact with a biochemicaltest sample. The materials and fabrication methods can be selected tooptimize high-volume manufacturing using plastic-based materials orhigh-sensitivity performance using semiconductor materials.

Representative examples of grating-type biosensors in the prior art aredisclosed in Cunningham, B. T., P. Li, B. Lin, and J. Pepper,Colorimetric resonant reflection as a direct biochemical assaytechnique. Sensors and Actuators B, 2002. 81: p. 316-328; Cunningham, B.T., J. Qiu, P. Li, J. Pepper, and B. Hugh, A plastic colorimetricresonant optical biosensor for multiparallel detection of label-freebiochemical interactions, Sensors and Actuators B, 2002. 85: p. 219-226;Haes, A. J. and R. P. V. Duyne, A Nanoscale Optical Biosensor:Sensitivity and Selectivity of an Approach Based on the LocalizedSurface Plasmon Resonance Spectroscopy of Triangular SilverNanoparticles. Journal of the American Chemical Society, 2002. 124: p.10596-10604.

The combined advantages of photonic crystal biosensors may not beexceeded by any other label-free biosensor technique. The development ofhighly sensitive, miniature, low cost, highly parallel biosensors andsimple, miniature, and rugged readout instrumentation will enablebiosensors to be applied in the fields of pharmaceutical discovery,diagnostic testing, environmental testing, and food safety inapplications that have not been economically feasible in the past.

In order to adapt a photonic bandgap device to perform as a biosensor,some portion of the structure must be in contact with a test sample.Biomolecules, cells, proteins, or other substances are introduced to theportion of the photonic crystal and adsorbed where the locally confinedelectromagnetic field intensity is greatest. As a result, the resonantcoupling of light into the crystal is modified, and thereflected/transmitted output (i.e., peak wavelength) is tuned, i.e.,shifted. The amount of shift in the reflected output is related to theamount of substance present on the sensor. The sensors are used inconjunction with an illumination and detection instrument that directslight into the sensor and captures the reflected or transmitted light.The reflected or transmitted light is fed to a spectrometer thatmeasures the shift in the peak wavelength.

The ability of photonic crystals to provide high quality factor (Q)resonant light coupling, high electromagnetic energy density, and tightoptical confinement can also be exploited to produce highly sensitivebiochemical sensors. Here, Q is a measure of the sharpness of the peakwavelength at the resonant frequency. Photonic crystal biosensors aredesigned to allow a test sample to penetrate the periodic lattice, andto tune the resonant optical coupling condition through modification ofthe surface dielectric constant of the crystal through the attachment ofbiomolecules or cells. Due to the high Q of the resonance, and thestrong interaction of coupled electromagnetic fields with surface-boundmaterials, several of the highest sensitivity biosensor devices reportedare derived from photonic crystals. See the Cunningham et al. paperscited previously. Such devices have demonstrated the capability fordetecting molecules with molecular weights less than 200 Daltons (Da)with high signal-to-noise margins, and for detecting individual cells.Because resonantly-coupled light within a photonic crystal can beeffectively spatially confined, a photonic crystal surface is capable ofsupporting large numbers of simultaneous biochemical assays in an arrayformat, where neighboring regions within ˜10 μm of each other can bemeasured independently. See Li, P., B. Lin, J. Gerstenmaier, and B. T.Cunningham, A new method for label-free imaging of biomolecularinteractions. Sensors and Actuators B, 2003.

There are many practical benefits for label-free biosensors based onphotonic crystal structures. Direct detection of biochemical andcellular binding without the use of a fluorophore, radioligand orsecondary reporter removes experimental uncertainty induced by theeffect of the label on molecular conformation, blocking of activebinding epitopes, steric hindrance, inaccessibility of the labelingsite, or the inability to find an appropriate label that functionsequivalently for all molecules in an experiment. Label-free detectionmethods greatly simplify the time and effort required for assaydevelopment, while removing experimental artifacts from quenching, shelflife, and background fluorescence. Compared to other label-free opticalbiosensors, photonic crystals are easily queried by simply illuminatingat normal incidence with a broadband light source (such as a light bulbor LED) and measuring shifts in the reflected color. The simpleexcitation/readout scheme enables low cost, miniature, robust systemsthat are suitable for use in laboratory instruments as well as portablehandheld systems for point-of-care medical diagnostics and environmentalmonitoring. Because the photonic crystal itself consumes no power, thedevices are easily embedded within a variety of liquid or gas samplingsystems, or deployed in the context of an optical network where a singleillumination/detection base station can track the status of thousands ofsensors within a building. While photonic crystal biosensors can befabricated using a wide variety of materials and methods, highsensitivity structures have been demonstrated using plastic-basedprocesses that can be performed on continuous sheets of film.Plastic-based designs and manufacturing methods will enable photoniccrystal biosensors to be used in applications where low cost/assay isrequired, that have not been previously economically feasible for otheroptical biosensors.

The assignee of the present invention has developed a photonic crystalbiosensor and associated detection instrument for label-free bindingdetection. The sensor and detection instrument are described in thepatent literature; see U.S. patent application publications U.S.2003/0027327; 2002/0127565, 2003/0059855 and 2003/0032039. Methods fordetection of a shift in the resonant peak wavelength are taught in U.S.Patent application publication 2003/0077660. The biosensors described inthese references include 1- and 2-dimensional periodic structuredsurfaces applied to a continuous sheet of plastic film or substrate. Thecrystal resonant wavelength is determined by measuring the peakreflectivity at normal incidence with a spectrometer to obtain awavelength resolution of 0.5 picometer. The resulting mass detectionsensitivity of <1 pg/mm² (obtained without 3-dimensional hydrogelsurface chemistry) has not been demonstrated by any other commerciallyavailable biosensor.

A fundamental advantage of the biosensor devices described in theabove-referenced patent applications is the ability to mass-manufacturewith plastic materials in continuous processes at a 1-2 feet/minuterate. Methods of mass production of the sensors are disclosed in U.S.Patent application publication 2003/0017581. As shown in FIG. 1, theperiodic surface structure of a biosensor 10 is fabricated from a lowrefractive index material 12 that is overcoated with a thin film ofhigher refractive index material 14. The low refractive index material12 is bonded to a base sheet of clear plastic material 16. The surfacestructure is replicated within a layer of cured epoxy 12 from asilicon-wafer “master” mold (i.e. a negative of the desired replicatedstructure) using a continuous-film process on a polyester substrate 16.The liquid epoxy 12 conforms to the shape of the master grating, and issubsequently cured by exposure to ultraviolet light. The cured epoxy 12preferentially adheres to the sheet 16, and is peeled away from thesilicon wafer. Sensor fabrication was completed by sputter deposition of120 nm titanium oxide (TiO₂) high index of refraction material 14 on thecured epoxy 12 grating surface. Following titanium oxide deposition,3×5-inch microplate sections are cut from the sensor sheet, and attachedto the bottoms of bottomless 96-well and 384-well microtiter plates withepoxy.

As shown in FIG. 2, the wells 20 defining the wells of the microtiterplate contain a liquid sample 22. The combination of the bottomlessmicroplate and the biosensor structure 10 is collectively shown asbiosensor apparatus 26. Using this approach, photonic crystal sensorsare mass produced on a square-yardage basis at very low cost.

The detection instrument for the photonic crystal biosensor is simple,inexpensive, low power, and robust. A schematic diagram of the system isshown in FIG. 2. In order to detect the reflected resonance, a whitelight source illuminates a ˜1 mm diameter region of the sensor surfacethrough a 100 micrometer diameter fiber optic 32 and a collimating lens34 at nominally normal incidence through the bottom of the microplate. Adetection fiber 36 is bundled with the illumination fiber 32 forgathering reflected light for analysis with a spectrometer 38. A seriesof 8 illumination/detection heads 40 are arranged in a linear fashion,so that reflection spectra are gathered from all 8 wells in a microplatecolumn at once. See FIG. 3. The microplate+biosensor 10 sits upon a X-Yaddressable motion stage (not shown in FIG. 2) so that each column ofwells in the microplate can be addressed in sequence. The instrumentmeasures all 96 wells in ˜15 seconds, limited by the rate of the motionstage. Further details on the construction of the system of FIGS. 2 and3 are set forth in the published U.S. Patent Application 2003/0059855.

The descriptions and discussions below refer to the label-freetechnology described above as BIND technology. BIND is a trademark ofthe assignee SRU Biosystems, Inc.

2. Fluorescence Amplification Sensors

U.S. Pat. No. 6,707,561 describes a grating-based biosensing technologythat is sometimes referred to in the art as Evanescent Resonance (ER)technology. This technology employs a sub-micron scale grating structureto amplify a luminescence signal (e.g., fluorescence,chemi-luminescence, electroluminescence, phosphorescence signal),following a binding event on the grating surface, where one of the boundmolecules carries a fluorescent label. ER technology enhances thesensitivity of fluorophore based assays enabling binding detection atanalyte concentrations significantly lower than non-amplified assays.

ER technology uses grating generated optical resonance to concentratelaser light on the grating surface where binding has taken place. Inpractice, a laser scanner sweeps the sensor at some angle of incidence(theta), typically from above the grating, while a detector detectsfluoresced light (at longer optical wavelength) from the sensor surface.By design, ER grating optical properties result in nearly 100%reflection, also known as resonance, at a specific angle of incidenceand laser wavelength (λ). Confinement of the laser light by and withinthe grating structure amplifies emission from fluorophores bound withinrange of the evanescent field (typically 1-2 um). Hence, at resonance,transmitted light intensity drops to near zero.

As noted above, the label-free biosensors described in theabove-referenced patent applications employ a sub-micron scale gratingstructure but typically with a significantly different grating geometryand objective as compared to gratings intended for ER use. In practicaluse, label-free and ER technologies have different requirements foroptical characteristics near resonance. The spectral width and locationof the resonance phenomena describes the primary difference. Resonancewidth refers to the full width at half maximum, in wavelength measure,of a resonance feature plotted as reflectance (or transmittance) versuswavelength (also referred to as Q factor above). Resonance width canalso refer to the width, in degrees, of a resonance feature plotted on acurve representing reflectance or transmittance as a function of theta,where theta is the angle of incident light.

Optimally, a label-free grating-based sensor produces as narrow aresonance peak as possible, to facilitate detection of small changes inpeak position indicating low binding events. A label-free sensor alsobenefits from a high grating surface area in order to bind morematerial. In current practice, one achieves higher surface area bymaking the grating deeper (though other approaches exist). Currentcommercial embodiments of label-free sensors produce a resonance near850 nm, thus BIND label-free detection instrumentation has beenoptimized to read this wavelength.

Conversely, practical ER grating sensor designs employ a relativelybroad resonance to ensure that resonance occurs at the fixed wavelengthlaser light and often fixed angle of incidence in the presence ofphysical variables such as material accumulation on the grating orvariation in sensor manufacture. Because field strength generallydecreases with resonance width, practical ER sensor design calls for abalance in resonance width. By choosing an appropriate ER resonancewidth, one ensures consistent amplification across a range of assay,instrument and sensor variables while maintaining ER signal gain. Atypical application uses a 633 nm wavelength to excite a popularfluorescent dye, known in the art as Cy5. Some ER scanninginstrumentation permits adjustments to incident angle to “tune” theresonance towards maximum laser fluorophore coupling. This practice,however, may induce an unacceptable source of variation without propercontrols.

Known ER designs also employ more shallow grating depths than optimallabel-free designs. For example, the above-referenced '561 patentspecifies the ratio of grating depth to “transparent layer” (i.e., highindex coating layer) thickness of less than 1 and more preferablybetween 0.3 and 0.7. Optimal label-free designs employ gratings with asimilarly defined ratio of greater than 1 and preferably greater than1.5. Label-free designs typically define grating depth in terms of thegrating line width or half period. For example, currently practicedcommercial label-free sensors have a half period of 275 nm and a gratingdepth of approximately 275 nm, thus describing a 1:1 geometric ratio.This same sensor design employs a high index of refraction oxide coatingon top of the grating with a thickness of approximately 90 nm. Thus,according to the definition in '561 patent, this sensor has a gratingdepth:oxide thickness ratio of approximately 3:1.

This disclosure reports grating-based sensor designs which areconstructed in a manner such that it is optimized for both modes ofdetection (label-free and fluorescence amplification), in a singledevice. Such a grating dramatically increases the diversity ofapplications made possible by a single product.

All the previously cited art is fully incorporated by reference herein.

SUMMARY

The following embodiments and aspects thereof are described andillustrated in conjunction with systems, tools and methods meant to beexemplary and illustrative, not limiting in scope. In variousembodiments one or more of the above-described problems have beenreduced or eliminated, while other embodiments are directed to otherimprovements.

In one aspect, a grating-based sensor is disclosed which is optimizedfor performance in both ER mode and in a label-free detection mode. Suchsensors exhibit a broad resonance at small angles of incidence (theta),mimicking the performance curves of a conventional ER grating biosensor,while also maintaining sharp resonance peak in a label-free detectionmode. Several representative embodiments are disclosed. A firstembodiment is optimized for ER mode in an air sample medium and with TMpolarization of light (perpendicular to the grating). A secondembodiment is disclosed which is optimized for ER mode in a liquidsample medium with 633 nm excitation, near normal incidence, and with TEpolarization. Computer modeling of both embodiments indicates that eachmaintains sharp peak wavelength resonance (high Q factor) in alabel-free detection mode.

In one configuration, a biosensor has a periodic surface gratingstructure (either in one or two dimensions), wherein the periodicgrating structure is constructed so as to optimize optical interrogationof the biosensor from a first light source in an evanescent resonance(ER) detection mode, and wherein the periodic grating structure isconstructed so as to optimize optical interrogation of the biosensorwith light from a second light source in a label-free detection mode. Inone possible embodiment, the grating takes the form of a two-dimensionalgrating, and wherein the grating is periodic in first and secondmutually orthogonal directions. In other embodiments, the grating is aone-dimensional grating, with periodicity in one direction (e.g., the Xdirection) but not in the second direction.

Label-free detection use of the biosensor benefits from deeper gratingsto provide more surface area, enabling more material attachment. Moreattached material generates more signal in the form of larger shift ofthe peak wavelength value. Prior ER gratings do not have enough surfacearea (depth) to render label-free sensitivity equivalent to currentlabel-free grating sensors. Hence, one designs biosensors maximizesurface area (translating greater grating depth in this case) whilemaintaining a broad resonance curve at the intended laser excitationwavelength and low angles of incidence, preferably less than 10 degreesand more preferably less than five degrees. Biosensors meeting ER andlabel-free performance requirements in representative one-dimensionalembodiments have a grating depth to half period ratio of between about0.6 and about 1.2. Grating depths of between approximately 160 nm andapproximately 210 nm are specifically contemplated. These parameters mayof course vary to address specific sensor performance objectives, toemphasizing ER or label-free performance, or for example intwo-dimensional gratings as disclosed herein.

Computer simulation of grating design, in accordance with the teachingsof this disclosure, will allow persons skilled in the art to developother grating designs in accordance with this disclosure which may varyfrom the specifics of the first and second embodiments and suchembodiments are offered by way of illustration and not limitation. In afurther aspect, methods of designing dual use ER and label-freedetection biosensors are disclosed using computer modeling techniques.

A grating-based sensor having a two-dimensional orthogonal gratingstructure suitable for both ER and label-free detection is alsodisclosed and may be preferred in some implementations. Atwo-dimensional grating can look like a waffle (holes), a waffle iron(posts), or a chessboard configuration with alternating high and lowregions in two dimensions. Two dimensional gratings can have differentperiods in the X and Y directions. These features may have variousprofiles in the Z direction such as angled or curved sidewalls. Thus, inthe case of the waffle pattern, the impressions or wells may have arectangular rather than a square shape. In practice, these features willalso appear rounded in the X and Y dimensions, i.e., will not have sharpcorners. Thus, the use of the terms “rectangular” and “square” areintended to refer to the overall configuration and allow for roundedcorners. This added flexibility provided by two dimensional gratingsallows one to tune the resonance positions for both label-free detectionand ER detection to occur at different wavelengths. This capabilityoffers significant benefit in terms of tuning of the ER resonance todifferent excitation wavelengths while maintaining compatibility withexisting label-free detection instrumentation. As an example, the Xperiodicity can provide a broad resonance at or near normal incidencewith wavelength tuned to excite the Cy3 fluorophore (green light) or theCy5 fluorophore (red light), while the Y periodicity can yield a sharplabel-free resonance between 820 and 850 nm (near infra red) similar tocurrently commercialized label-free sensors.

Two-dimensional, two-level grating structures are also disclosed as afurther embodiment of a grating-based biosensor which is structured andarranged to have good performance for both ER and label-free detection.

In another aspect, a method of analyzing at least one sample isdisclosed comprising the steps of placing the at least one sample on abiosensor comprising a substrate having a periodic surface gratingstructure, wherein the periodic grating structure is constructed anddesigned for optical interrogation of the biosensor in an evanescentreflection (ER) detection mode as well as optimize optical interrogationof the biosensor in a label-free detection mode. The method furthercomprises the steps of illuminating the biosensor in a readout detectioninstrument with light from a light source designed for the ER detectionmode and illuminating the biosensor with light from the light source (orpossibly from a second light source) designed for the label-freedetection mode; and analyzing light reflection from the biosensor. Theanalyzing of the sample may include detecting binding of a component ofthe sample, e.g., binding of the component of the sample to the surfaceof the biosensor or binding of second sample component (e.g.,fluorophore, inhibitor or label) to a first sample component (e.g.,protein)

In one embodiment, the readout system includes two light sources, onefor BIND (e.g., a while light source or light emitting diode) and asecond light source such as laser for ER measurements. However, in otherembodiments a single light source is provided such as a Xenon dischargelamp or tunable laser, with two (or more) bandpass filters sampling thelight source to provide appropriate illumination wavelengths for the twosensing modes.

In one possible embodiment, the sample is in an air medium, and whereinthe light from the first light source has a polarization perpendicularto the grating structure. In another possible embodiment, the sample isin a liquid medium, and wherein the light from the first light sourcehas a polarization parallel to the grating structure. In one possibleembodiment, light from the first light source has a wavelength selectedto activate a fluorophore bound to the sample. In another possibleembodiment, light from the first light source has a wavelength selectedto activate native fluorescence of the sample. In still otherembodiments, a fraction of the sample is bound to an inhibitor, whichmay include a bound fluorophore. The sample may be, for example, aprotein.

Several representative configurations of a readout and detectioninstrument for the inventive biosensor are also disclosed. In oneembodiment, the readout and detection instrument includes a first lightsource adapted for obtaining ER data from the biosensor; a second lightsource adapted for obtaining label-free detection data; an opticalsystem combining the light from the first and second light sources intoan illuminating beam for illuminating the biosensor; at least onedetector for detecting reflected light from the biosensor; and ananalysis module using data from the at least one detector and obtainingER and label-free data from the sample. The detector may be an imagingdetector such as a charge-coupled device (CCD imager). Other types ofdetectors are also envisioned, such as photodetector, spectrometer, or acombination thereof, one for acquiring ER data and one for acquiringBIND data. In another representative configuration, the optical systemselectively illuminates the biosensor with light from a single lightsource. The biosensor may have multiple detection sites or wells, andthe instrument may include a motion stage for successively moving thedetection sites relative to the light sources to sequentially obtain ERand label-free data from all the detection sites.

In sum, this disclosure describes a novel detection and quantificationplatform that combines a photonic crystal based label-free biosensorwith enhanced fluorescence capabilities, in a single device. Alone,label-free and ER technologies have great utility. The ability to jointhese two detection technologies in a single biosensor creates apowerful approach for universal detection and selective measurement ofinteraction between and within biological materials such as cells,proteins, and small molecules. The combined biosensor of this disclosureis useful for detection of a broad range of biological or chemicalsample entities. Examples of the types of samples which can be detectedinclude small and smaller molecular weight molecules (i.e., substancesof molecular weight<1000 Da and between 1000 Da to 10,000 Da), aminoacids, nucleic acids, lipids, carbohydrates, nucleic acid polymers,viral particles, viral components and cellular components such as butnot limited to vesicles, mitochondria, membranes, structural features,periplasm, or any extracts thereof.

In general, further examples of specific binding substances (samples)which may be detected with the biosensor of this invention includepolypeptides, antigens, polyclonal antibodies, monoclonal antibodies,single chain antibodies (scFv), F(ab) fragments, F(ab′)2 fragments, Fvfragments, small organic molecules, cells, viruses, bacteria, polymers,peptide solutions, protein solutions, chemical compound librarysolutions, single-stranded DNA solutions, double stranded DNA solutions,combinations of single and double stranded DNA solutions, RNA solutionsand biological samples. Such biological samples could consists of, forexample, blood, plasma, serum, gastrointestinal secretions, homogenatesof tissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid,amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavagefluid, semen, lymphatic fluid, tears and prostatic fluid.

The biosensor described herein may be used to detect (a) binding ofcomponents any of these types of samples to the biosensor surface, (b)binding of the sample to another component of the sample, e.g., afluorophore in the sample, and (c) binding of the sample or samplecomponent to a second sample which is added to the sample. As an exampleof binding (b), the sensor surface may bind to some component of thesample, such as for example streptavidin-biotin or 6His, and thebiosensor may be used to detect the interaction of the bound componentof the sample with an additional grouping of components in the sample,such as a polymerase complex. In the latter example of binding (c), asample may have a component that is attached to the surface of thebiosensor and another component which specifically binds/attractsanother component(s) from a second sample that is placed on thebiosensor.

The sensor of this disclosure may also be used for quantification of theamount of material binding or interaction.

The following general examples by no means represent a complete orexclusive listing of novel utilities enabled by such a dual usegrating-based sensor as disclosed herein:

1. Combined, the two technologies distinguish the percentage offluorophore labeled material present in a mixed sample population. Thelabel-free signal provides a quantitative measure of the total massbound to the sensor while the ER signal quantifies the presence of thelabel.

2. The combination can also increase statistical rigor in themeasurement of interactions between and within cells, proteins, andsmall molecules by providing duplicate binding signals from differentsources.

3. Utilizing the Förster Resonant Energy Transfer (FRET) principle, thedual use sensor may enable measurement of the distance between twodifferentially labeled fluorescent molecules or two differentiallylabeled portions of the same molecule. The label-free signal quantifiesmolecular density.

4. The combination of the two technologies can provide additiveinformation. The label-less signal can quantify the attachment of cellswith the fluorescent signal quantifying the amount of fluorophorelabeled ligand bound to the cell. Other scenarios are of course possiblewhere label-less and labeled biological entities, such as those listedabove, are detected on the inventive biosensor.

5. The combination of the two technologies may provide a measure of themolecular mass by distinguishing molecular count from total bound mass.

6. The combined biosensor further permits two different independentquantification tests to be performed for other scenarios such as thestudy of inhibition binding. Furthermore, a more complete understandingand characterization of inhibition binding interactions between aprotein and a substrate is possible, including the ability to directlyquantify inhibitor ligand binding. As an additional example, thebiosensor facilitates the study of very tight binding interactionswhereby a known competitive inhibitor with a weaker binding affinity isemployed to perturb/observe the much tighter binding entity.

7. The combined ER and label-free biosensor is particularly useful forassays which utilize the natural fluorescence of biological molecules(i.e., without requiring the use of a bound fluorescence label), to makebiophysical characterization measurements of activity such as folding,stacking, and changes and rates of changes to these upon interactionswith other biological molecules and small test molecules. Suchcharacterization measurements could be made using a bound fluorescencelabel, but such bound label is not necessarily required, especially forbiological materials having an inherent fluorescence property. SeeCharles R. Cantor and Paul R. Schimmel, parts 1-3 BiophysicalChemistry—The behavior and study of biological molecules, W.H. Freemanand Company, New York, (1980), page 443 and table 8-2 for a listing offluorescence characteristics of protein and nucleic acid constituentsand coenzymes, their absorption and emission spectra and sensitivity.This technique of using native fluorescence is especially important withnucleic acid polymers (DNA, RNA) (fluorescent nucleoside bases) stackingand hybridization, proteins (fluorescent amino acids phenylalanine,tryptophan, and tyrosine) and lipid membranes (enhancement and quenchingeffects upon incorporation of fluorophores into their differentcompartmentalizations). In one embodiment, the label-free BIND featureallows the quantification of the amount of sample material or ligandbound thereto and the ER feature detects the native fluorescence andallows the sensitive tracking of the biophysical change.

BRIEF DESCRIPTION OF THE DRAWINGS

Exemplary embodiments are illustrated in referenced figures of thedrawings. The embodiments and figures disclosed herein are offered byway of example and not limitation. All questions regarding scope of theinvention are to be answered by reference to the claims.

FIG. 1 is an illustration of a prior art biosensor arrangement.

FIG. 2 is an illustration of a prior art biosensor and detection systemfor illuminating the biosensor and measuring shifts in the peakwavelength of reflected light from the biosensor.

FIG. 3 is an illustration of an arrangement of 8 illumination heads thatread an entire row of wells of a biosensor device comprising thestructure of FIG. 1 affixed to the bottom of bottomless microtiterplate.

FIG. 4 is a cross-section of a first embodiment of a combined ER andlabel-free detection biosensor.

FIG. 5 is a cross-section of a second embodiment of a combined ER andlabel-free detection biosensor.

FIG. 6 is a graph comparing transmission as a function of incident angletheta for a prior art ER biosensor (“NovaChip”) with a computersimulation of the embodiment of FIG. 4 when used for ER detection in adry (air) medium environment.

FIG. 7 is a graph comparing reflection as a function of wavelength forthe embodiment of FIG. 4 in a label-free detection mode in an aqueousmedium environment; the graph of FIG. 7 was generated from a computersimulation of the embodiment of FIG. 4.

FIG. 8 is a graph of reflection as a function of wavelength for theembodiment of FIG. 4 in a label-free detection mode, showing a shift ofthe peak wavelength value in response to surface mass addition (e.g., byadding a sample to the biosensor). The graph of FIG. 7 was alsogenerated via a computer simulation of the embodiment of FIG. 4.

FIG. 9 is a graph of reflection on as a function of wavelength showingthe resonance peaks for the embodiment of FIG. 5 in label-free and ERdetection modes. The graph was generated from a computer simulation ofthe embodiment of FIG. 5.

FIG. 10 is a graph of transmission as a function of theta for theembodiment of FIG. 5 and comparing the curve with the transmission curveof a prior art “NovaChip” example of an ER sensor.

FIGS. 11A and 11B are perspective and cross-sectional views,respectively, of a one-dimensional linear grating structure designedsolely for ER detection, modeled as a rough approximation of an ER chipdisclosed in a prior art article of Budach et al.

FIGS. 12A and 12B are graphs of the reflection efficiency as a functionof wavelength and incidence angle, respectively, obtained when lightpolarized in the X direction is incident on the structure of FIGS. 11Aand 11B.

FIGS. 13A-13C are plots of the X, Y and Z components of electric fieldamplitude in the XY plane corresponding to the lower surface of thestructure of FIGS. 11A and 11B located a Z=110 nm, for incidentwavelength 632 nm. FIGS. 13D-13F plot of the X, Y and Z components ofthe magnetic field amplitude in the same XY plane represented in FIGS.13A-13C for incident wavelength 632 nm.

FIGS. 14A-14C are plots of the X, Y and Z components of electric fieldamplitude in the XY plane corresponding to the upper surface of thestructure of FIGS. 11A and 11B located at Z=140 nm, for incidentwavelength 632 nm. FIGS. 14D-14F plot the X, Y and Z components of themagnetic field amplitude in the same XY plane represented in FIGS.14A-14C for incident wavelength 632 nm.

FIGS. 15A and 15B are perspective and cross-sectional views,respectively, of a two-dimensional grating design characterized byperiodic holes in a grating structure which is optimized for BIND(label-free) detection in a water environment when illuminated by Xpolarized light and optimized for ER detection in an air environmentwhen illuminated by Y polarized light.

FIGS. 16 and 17 are graphs of the reflection efficiency as a function ofwavelength and incidence angle (632.5 nm), respectively, obtained whenlight polarized in the Y direction is incident on the structure of FIGS.15A and 15B. These figures demonstrate utility in the ER mode.

FIGS. 18A-18C are plots of the X, Y and Z components of electric fieldamplitude in the XY plane corresponding to the lower surface of thestructure of FIGS. 15A and 15B at Z=78 nm, for incident wavelength 632.5nm. FIGS. 18D-18F plot the X, Y and Z components of the magnetic fieldamplitude in the same XY plane represented in FIGS. 18A-18C for incidentwavelength 632.5 nm.

FIGS. 19A-19C are plots of the X, Y and Z components of the electricfield amplitude in the XY plane corresponding to the upper surface ofthe structure of FIGS. 15A and 15B at Z=433 nm for incident wavelength632.5 nm. FIGS. 19D-19F are plots of X, Y and Z components of themagnetic field amplitude in the same XY plane represented in FIGS.19A-19C for incident wavelength 632 nm.

FIG. 19G is a graph of reflection efficiency as a function of wavelengthfor the embodiment of FIG. 15 obtained when illuminated by lightpolarized in the X direction. This resonance peak is used for label-freedetection.

FIGS. 20A and 20B show perspective and cross-sectional views,respectively, of a two-dimensional grating design characterized byperiodic posts in a grating structure which is optimized in onedirection for BIND (label-free) detection in a water environment whenilluminated by X polarized light and optimized for ER detection in anair environment when illuminated by Y polarized light.

FIGS. 21A and 21B graph the reflection efficiency as a function ofwavelength and incidence angle (633 nm wavelength), respectively, whenlight polarized in the X direction is incident on the structure of FIGS.20A and 20B. The figures demonstrate utility in the ER mode.

FIGS. 22A-22C are plots of the X, Y and Z components of electric fieldamplitude in the XY plane corresponding to the lower surface of thestructure of FIGS. 20A and 20B at Z=70 nm, for incident wavelength 633nm. FIGS. 22D-22F plot the X, Y and Z components of the magnetic fieldamplitude in the same XY plane represented in FIGS. 22A-22C for incidentwavelength 633 nm.

FIGS. 23A-23C plot the X, Y and Z components of the electric fieldamplitude corresponding to the upper surface of the structure of FIGS.20A and 20B at Z=430 nm for incident wavelength 633 nm. FIGS. 23D-23Fplot the X, Y and Z components of the magnetic field amplitude in thesame XY plane represented by FIGS. 23A-C for incident wavelength 632 nm.

FIG. 24 is a graph of reflection efficiency as a function of wavelengthfor the embodiment of FIG. 20 obtained when illuminated by lightpolarized in the X direction. This resonance peak is used for label-freedetection.

FIG. 25 is a schematic drawing of an imaging readout system for acombined ER and label-free grating-based sensor.

FIG. 26 is a schematic illustration of a second readout system for acombined ER and label-free grating-based sensor.

FIG. 27 is a more detailed illustration of the embodiment of FIG. 26.

FIGS. 28A-C are three views of a unit cell showing a two-level,two-dimensional grating structure for yet another embodiment of acombined ER and label-free sensor.

FIG. 29 is a graph of the reflection spectrum (relative intensity as afunction of reflected light wavelength) obtained using a computersimulation of the structure of FIG. 28A-C.

FIG. 30 is a cross-sectional view of a combined ER and BINDgrating-based sensor in which an intermediate SiO2 layer is addedbetween the UV-cured plastic grating layer and the high index ofrefraction layer forming the upper surface of the sensor.

FIG. 31 is an image of a microarray of spots deposited on agrating-based sensor (which may or may not be optimized for both ER andBIND measurements).

FIG. 32 is a graph of a peak shift for one of the spots of FIG. 31 dueto presence of a DNA sample being placed on the sensor.

FIG. 33 is an illustration of a row of spots showing the shift in peakwavelength value in nanometers (which is quantitatively related to theamount of DNA in the spot) as a function of position and showing amissing spot (dark location in the row of spots), the missing spotindicated by the low region in the graph.

DETAILED DESCRIPTION

Grating-based biosensors are disclosed which have a periodic gratingconstruction which is optimized and useful for both ER detection, eitherin a liquid or dry environment, and for label-free detection. A readoutsystem adapted for use with the biosensors is also described. Methods oftesting a sample with the inventive biosensors are also described.

First Embodiment

FIG. 4 is a schematic cross-sectional illustration of a first embodimentof a one-dimensional sensor having a grating structure 100 that isexpected to meet commercial requirements for both ER and label-freeapplications of a grating-based sensor. FIG. 4 shows one period of agrating structure 100 in one dimension or direction. The dimensions arenot to scale in FIG. 4.

The grating 100 of FIG. 4 is superimposed and bonded to a base sheet ofclear material such as Polyethylene Terepthalate (PET) or other plastic,glass or other material (not shown).

The grating structure consists of a periodically repeating material 102which preferably comprises a UV-cured material, e.g., epoxy, appliedwith the aid of a grating master wafer (not shown) to replicate thegrating pattern onto the base sheet of PET material located below thelayer “substrate.” The UV cured material 102 is applied to a substratesheet such as PET. Substrate materials can also include polycarbonate orcyclo-olefin polymers such as Zeanor®. Other means of producing thestructured layer 102 include thermally stamping directly into a polymersubstrate. The middle material 104 represents a sputtered oxide coatingwith high refractive index (e.g. TiO₂ or Ta₂O₅). The upper most material106 represents a medium for a sample, which is normally either awater-based buffer, for label-free detection mode, or air, for ER mode.The structure has the periodicity, layer structure, and horizontaltransition points as shown in the Figure. The specifics of the design ofcourse may change while still providing good performance for bothlabel-free detection and ER detection.

The design of FIG. 4 was developed and its performance modeled with theaid of a computer and a software program GSolver (Grating SolverDevelopment Co., Allen Tex., www.gsolver.com). The various geometricaldimensions and parameters, spacing, well depth, materials, and index ofrefraction data associated with the materials allows the design to bestudied on a computer and simulations run to predict the Transmission v.Theta curve and reflection as a function of wavelength curve. Suchsimulations can be run in situations where the sample is dry and whenthe sample is suspended in water or other fluid medium with known indexof refraction. Such simulations allow the designer to optimize, i.e.,change, the various design parameters (thicknesses, transitions, period,etc.) to satisfy the requirements for both ER and label-free detection.

ER technology heretofore employs a resonance mode induced by incidentlight with a polarization parallel to the grating, defined here as TEmode or polarization. Label-free detection technology typically employsa resonance mode induced by incident light with polarizationperpendicular to the grating, defined here as the TM mode orpolarization. This mode produces the narrowest resonance when the sampleis suspended in a liquid medium.

In the first embodiment of FIG. 4, a grating biosensor design isdescribed which utilizes TM polarization for both label-free detectionof a sample suspended in liquid and ER detection in an air (dry)environment. Changing the medium above the grating from water to airresults in a change in resonance characteristics from those useful forlabel-free detection to those useful for ER amplification of dyesresponding to 633 nm excitation. The design of FIG. 4 is notspecifically optimized to ER detection in a water mode and may not evenwork acceptably for ER in a water mode. However, many ER detectionassays are run in an air environment and so the design of FIG. 4 hasmuch utility for ER detection.

FIG. 6 is a graph that compares Transmission v. Theta data for a priorart ER device (NovaChip, Novartis AG) with a computer simulation ormodel of the first design of FIG. 4 (“Combind 400 Air”). NovaChip datais disclosed in Budach et al., Generation of Transducers forFluorescence-Based Microarrays with Enhanced Sensitivity and TheirApplication to Gene Expression Profiling, Analytical Chemistry (2003)and in Neuschafer et al., Evanescent resonator chips: a universalplatform with superior sensitivity for fluorescence-based microarrays,Biosensors and Bioelectronics 18 (2003) 489-497. The curves 110 and 112have a similar shape suggesting the simulated device (ComBIND 400) wouldfunction equivalently to the ER device. The NovaChip TE resonance occursat ˜2 degrees from normal incidence. The first design of FIG. 4 producesTE resonance at ˜3 degrees, which is considered only a minor differencegiven that one can adjust the angle of incident light (see thediscussion of the readout and detection instrument for the sensor laterin this disclosure).

The graph of FIG. 7 plots simulated reflection vs. wavelength for theComBind 400 design (FIG. 4). The broad reflection peak centered around628 nm corresponds to the ER-air mode resonance occurring at ˜3 degreesin the Transmission v. Theta curve above. The narrow peak, labeled“water”, serves for the label-free mode of detection. Note the extremesharpness of the peak 114. This suggests that the design of FIG. 4 wouldwork well for label-free detection in a water environment.

During label-free mode detection, biological molecules adhere to theTiO₂ coating and effectively increase the optical thickness of thatmaterial. This results in a shift in the peak wavelength value (PWV) ofthe resonance. A larger PWV shift for a fixed amount of materialrepresents higher detection sensitivity. When comparing grating designsin a computer simulation, the simulation of additional biologicalmaterial can be modeled by incrementing the thickness of the TiO₂ layerrather than adding a hypothetical biological layer. This method hasproven effective in other grating design exercises.

FIG. 8 is a graph that plots the peak wavelength value in a waterenvironment before and after the addition of a certain amount ofsimulated mass (simulated by increasing the thickness of the TiO₂ layer104 of FIG. 4). The peak position shifts to higher wavelength, as isexpected in label-free biosensor operation. The ratio of wavelengthshift to simulated mass is equivalent to that of the commercializedbiosensors of the applicant's assignee. Hence, the grating of FIG. 4 isexpected to yield equivalent label-free performance to the currentlabel-free biosensor gratings.

To summarize, simulations predict dual-use capabilities for the gratingdesign disclosed in FIG. 4. When dry, it can amplify fluorescent bindingsignals according to the technology known as evanescent resonance (ER).When wet, the grating performs as well as a label-free detectoraccording to the technology known as guided mode resonance detection orcommercially as BIND (trademark of SRU Biosystems, Inc.), available fromthe applicants' assignee SRU Biosystems, Inc.

Second Embodiment

FIG. 5 is a cross-section of a second embodiment, showing one period ofthe grating structure in one dimension and the structure of the of theUV cured layer 102, the high index of refraction layer 104, and thesample medium 106. The dimensions and transition points are as shown inthe drawing. The drawing is not to scale.

The design of FIG. 5 differs from that of FIG. 4 is several respects:

a) It has a shorter grating period.

b) It has narrower grating troughs or recesses. The “duty cycle”(percentage of the grating at the upper level in a unit cell) is 88% inFIG. 5 (0 to 0.85 and 0.97 to 1.0). Narrow troughs with duty cycles ofbetween 70 and 95% are exemplary of the narrow trough embodiments. Thenarrow troughs generally give better label-free detection results. Thenarrow trough feature narrows the TE resonance peak, thus indicatingincreased field strength. While practical use of the ER effect requiresa sufficiently broad resonance, a resonance with excessive width willhave insufficient field strength to produce useful fluorescence signalamplification

c) It has a 1:1 ratio of grating depth to half period.

The design of FIG. 5 exemplifies a one-dimensional sensor that enablesboth label-free and ER operation in a water (or buffer) environment.This contrasts with the design of FIG. 4, which is designed for ERoperation in air and label-free operation in water. The graph of FIG. 9shows the ER (TE polarization) and comparatively narrow label-free (TMpolarization) spectral resonance characteristics in a water environmentfor the design of FIG. 5. The graph was generated from a computersimulation of the embodiment of FIG. 5. The graph of FIG. 10 comparesthe TE angular resonance, using 633 nm excitation, of the FIG. 5 design(“ComBIND 370”) as compared to a prior art NovaChip. In this case, thesimulated transmission minimum 116 occurs below five degrees angle ofincidence, close to that of the existing NovaChip ER device. Theincident angle, period, excitation wavelength, high index of refractionmaterial thickness, grating duty cycle, and grating depth allinterrelate. Excitation wavelengths for commercial fluorophores areknown and can be looked up. Angles of incidence of less than 25 degreesshould be acceptable but angles near normal (Theta close to zero) arepreferred. With angle and wavelength confined to narrow ranges,designing a grating with functional and commercially useful ER andlabel-free performance one must determine a grating period, duty cycle,depth and high index of refraction material thickness that result inhigh PWV shift in response to mass attachment, for label-free use, andhigh surface field at the excitation wavelength of the specifiedfluorophore or dye for ER mode. Additionally, the design must produce alabel-free resonance with spectral width as narrow as possible whilemaintaining an ER resonance with angular width enough to yield apractical parameter window for measurement. The design may incorporateperformance trade-offs. For example, optimization of ER performanceengages a trade-off between field strength, which yields signalamplification, and tolerance for sensor, instrument and assay variables.Narrower ER resonance generally indicates higher field strength whilebroader ER resonance provides increased measurement tolerance.Typically, label-free and ER performance optimization involves anothertrade-off between grating depth, which enhances label-free performance,and ER resonance width. For example, in the case of the design of FIG.5, increasing the grating depth widens the TE/ER resonance beyondoptimum. Increasing the duty cycle (narrowing the troughs) compensates,narrowing the resonance back towards optimum, and thus maintaining fieldER strength.

Thus, one preferred approach to finding a dual use granting structurefor both ER and label-free detection modes involves finding a gratingwith depth to half period ratio approximately in the range of 0.6 to 1.2or more that also yields a broad angular resonance in either TM mode inan air environment, TE mode in air environment, or TE mode in a waterenvironment, at the excitation wavelength of interest (e.g., 633 nm) anda resonance angle less than 25 degrees. This broad resonance preferablyhas a width between 1 degree and 10 degrees or, in terms of spectralwidth, between 5 nm and 30 nm. Such design efforts can be readilyimplemented in a computer, e.g., using the Gsolver software. Moredirectly, one can comparatively model field strength at the gratingsurface using a software such as R-Soft, available from RSoft DesignGroup, www.rsoftdesigngroup.com.

Grating depths in the range of 100 to 600 nm and grating periods in therange of 300 to 600 nm are considered exemplary.

Persons skilled in the art having the benefit of this disclosure will beable to model potential grating designs on a computer and arrive atsuitable designs in accordance with this invention.

Two-Dimensional Gratings

The possibility of a two-dimensional (2-D) grating structure, suitablefor both ER and label-free detection, is also contemplated and may bepreferred. A two-dimensional grating can look like a waffle (holes), awaffle iron (posts), or a chessboard configuration with alternating highand low regions in two dimensions. Two-dimensional gratings can havedifferent periods in the X and Y directions. These features may havevarious profiles in the Z direction such as angled or curved sidewalls.Thus, in the case of the waffle pattern, the impressions or wells mayhave a rectangular rather than a square shape. This added flexibilityallows one to tune the resonance positions for both label-free detectionand ER detection to occur at different wavelengths. This flexibilityoffers significant benefit in terms of tuning the ER resonance todifferent excitation wavelengths while maintaining compatibility withexisting label-free detection instrumentation. As an example, the Xperiodicity can provide a resonance at or near normal incidence withwavelength tuned to excite the CY3 fluorophore (green light) or the CY5fluorophore (red light), while the Y periodicity can yield a resonancefixed between 820 and 850 nm (in the near infra red).

The examples of 2-D biosensor structures described herein were developedusing computer simulations and Rigorous Coupled Wave Analysis (RCWA)with a commercially available software package (RSoft). The computersimulations enable the device designer to vary the physical parametersof the device (refractive index, thickness, width, height, structuralshape) to determine: 1) the electromagnetic field distribution withinand around the device, 2) the reflectance or transmittance behavior as afunction of incident angle of light and wavelength of light, and 3) howthe reflected (or transmitted) spectrum is changed by the attachment ofbiomolecular material to the surface of the biosensor.

The specific 2-D embodiments described herein are optimized for combineddetection by BIND and ER methods in a single device where the sensorcontacts water during the BIND measurement and air during the ERmeasurement. Any combination of dry and wet for BIND and ER may besimilarly optimized (e.g., measure both BIND and ER in a wet mode).

To more fully appreciate the advantages provided by the combined ER andBIND (label-free) two-dimensional device, a discussion will initially bepresented in conjunction with FIGS. 11-14 of a linear (one-dimensional)structure optimized for ER only. Simulations were first performed on theER-only structure. The structure corresponds approximately to a priorart ER chip published by Budach et al., Generation of Transducers forFluorescence-Based Microarrays with Enhanced Sensitivity and TheirApplication for Gene Expression Profiling, Anal Chem 2003, 75,2571-2577. (Note: The Budach et al. grating is a linear grating and sois a 1-D structure as that term is used herein. The thinner TiO₂ highindex of refraction of material of FIG. 11A-11B as compared to thethicker Ta₂O₅ layer described by Budach et al. paper achieves a deviceof equivalent optical “thickness” by taking into account the differentindices of refraction of the two materials. The modeling of FIG. 11A-11Bis not meant to exactly replicate the Budach et al. device, but ratherto approximate it.)

The simulations were done to determine the electromagnetic fielddistribution as well as reflection as a function of angle and wavelengthfor a representative device for ER enhancement of Cy5 dye. Inparticular, FIGS. 11A and 11B are perspective and cross-sectional views,respectively, of a one dimensional linear grating structure designedsolely for ER detection. As shown, the structure has a linear gratingprofile consisting of a 30 nm raised ridge 201 and a 110 nm TiO₂ layer203 covering the raised ridge 201. The periodicity is in the Y-direction(ridge 201 repeating every 356 nm.

FIGS. 12A and 12B graphs the reflection efficiency as a function ofwavelength and incidence angle, respectively, of the structure of FIGS.11A and 11B, as determined by RCWA. Note that at normal incidence, thepeak wavelength (632 nm in FIG. 12A) corresponds to the excitationwavelength of Cy5, and that there is a broad range of angles (FIG. 12B)with high reflection efficiency of the 632 nm wavelength when theincident light is rotated at an angle, theta, in a parallel directionwith respect to the grating line or ridge 201.

FIGS. 13A-13C are plots of the X, Y and Z components of electric fieldintensity in the XY plane corresponding to the lower surface of thestructure of FIGS. 10A and 10B located a Z=110 nm, for incidentwavelength at 632 nm. FIGS. 13D-13F plot of the X, Y and Z components ofthe magnetic field intensity in the same XY plane represented in FIGS.13A-13C for incident wavelength at 632 nm.

The plots of FIG. 13 show the strength of the three components of theelectric field vector (Ex, Ey, and Ez) and the magnetic field vector(Hx, Hy, and Hz) as a function of XY position on the lower exposedsurface of the device. The upper exposed portion of the structure 200 isshaded here because the upper surface lies in a different horizontalplane than the lower surface. In the computer simulation, the sensor isilluminated with a light source having a 1 V/m magnitude electric fieldand a 1 A/m magnetic field at the resonance wavelength. Hence, fieldstrength values greater than 1 represent concentration of fieldintensity at the sensor surface resulting from resonance. The power ofthe electromagnetic field is calculated by the cross product of E and Hfield components. The field power, at a given location on thestructure's surface, specifies the energy available to excitefluorophores bound to the structure's surface. Higher power will, intheory, result in higher fluorescence emission. The plots show that theelectric and magnetic fields, and thus the power, do not distributeevenly over the structure's surface, but instead locations exist withhigher than average power (areas in red and orange in a color version ofthe drawing indicated at 202) and with lower than average power (areasin a color version of the drawing, areas in violet or blue, indicated at204).

FIGS. 14A-14C are plots of the X, Y and Z components of electric fieldintensity in the XY plane corresponding to the upper surface of thestructure of FIGS. 11A and 11B located at Z=140 nm, for incidentwavelength 632 nm. FIGS. 14D-14F plot the X, Y and Z components of themagnetic field intensity in the same XY plane represented in FIGS.14A-14C for incident wavelength 632 nm.

These plots (FIG. 14) show the strength of the three components of theelectric field vector (Ex, Ey, and Ez) and the magnetic field vector(Hx, Hy, and Hz) as a function of XY position on the upper exposedsurface of the device. The lower exposed portion of the structure isshaded here as indicated at 200 because the upper surface lies in adifferent horizontal plane than the lower surface. As with the plots ofFIG. 13, in the simulation the sensor is illuminated with a light sourcehaving 1 V/m electric field, 1 A/m magnetic field amplitudes at theresonant wavelength. Hence, field strength values greater than 1represent the concentration of field intensity at the sensor surface,resulting from resonance. The cross products of E and H fieldcomponents, as before, represent the instantaneous power distribution,at the resonant wavelength, available for fluorophore excitation.

A. Holes Embodiment Example

Now, a specific example of a 2D “holes” embodiment of a combinedbiosensor will be described in conjunction with FIGS. 15-19. Thebiosensor is constructed in two dimensions so as to be optimized forboth ER and label-free (BIND) detection using a single device.

FIGS. 15A and 15B provide perspective and cross-sectional views,respectively, of a unit cell for a two-dimensional grating designcharacterized by periodic holes 210 in a grating structure. The gratingdesign optimizes for water mode BIND (label-free) detection and air modeER detection. The device includes an upper TiO₂ layer 104 of 78nmthickness and a lower substrate 102 layer of UV-cured material having agrating pattern as shown applied to a base substrate sheet.

The two-dimensional unit cell shown in FIGS. 15A and 15B differentiatesfrom the one-dimensional linear grating design of FIGS. 11A and 11B. Thestructure of FIGS. 15A and 15B is designed in such a way that incidentlight polarized perpendicular to the X-axis, as shown, produces a BINDsignal, incident light polarized perpendicular to the Y-axis enables ERmeasurement. Using this design method, the BIND and ER resonantwavelengths (at a particular angle of incidence—preferably near normalincidence) may be chosen independently, and so the respective BIND andER resonant wavelengths may occur at very different values. The combinedBIND/ER structure described in this embodiment is optimized to provide aBIND resonance in the near infrared (˜800-900 nm) wavelength region,while providing an ER resonance at 632.5 nm for excitation of the Cy5fluorophore. In this example, the design assumes a water environmentover the sensor during BIND measurement and an air environment over thesensor during ER measurement. The differing wavelength requirements forER and BIND engender selection of a unit cell with a rectangular “hole”(210). Thus, the unit cell may have differing dimensions in the X and Ydirections. For example, the period in the X direction is 550 nm for theBIND wavelength, but is 432 nm in the Y direction as required for thelower wavelength ER resonance. The fabrication process dictates that thehigh refractive index dielectric thickness will be the same in the X andY directions. For fabrication simplicity, the design also has uniformgrating depth. The fabrication process will also result in rounding ofthe hole corners, however the principal function of the design remainsunchanged. One skilled in the art will appreciate that when a computeris used to generate and test a design such as shown in FIGS. 15A and15B, the designer can change the specific dimensions of the unit cell,grating depth, and coating layers and run simulations of fieldintensity, peak wavelength, reflectance as a function of theta, andother tests and may select other dimensions while still achievingacceptable results. Thus, the example of FIGS. 15A and 15B is meant tobe an illustrative embodiment and not limiting in scope.

FIGS. 16 and 17 are graphs of the reflection efficiency as a function ofwavelength and incidence angle, respectively, for the structuresdisclosed in FIGS. 15A and 15B when illuminated with light polarizedalong the Y axis. These figures, generated by RCWA, represent operationin an ER mode. FIG. 17 shows that, at the resonant wavelength, thereflected intensity as a function of incident angle observes theacceptance angle for light that will induce a significant ER effect. Inphysical measurements, the double peak and dip between the peaks in theplot of FIG. 16 may not resolve.

In a similar manner to the 1D example above, RCWA calculations may beused to determine the spatial distribution of the amplitude of theelectric field magnitude components (Ex, Ey, and Ez) and the magneticfield components (Hx, Hy, Hz) at the ER resonant wavelength for thestructure illustrated in FIGS. 15A and 15B. FIGS. 18A-18C plots the X, Yand Z components of electric field intensity in the XY planecorresponding to the lower surface of the structure of FIGS. 15A and 15Bat Z=78 nm, for incident wavelength 632.5 nm. FIGS. 18D-18F plot the X,Y and Z components of the magnetic field intensity in the same XY planerepresented in FIGS. 18A-18C for incident wavelength 632.5 nm. Thesefield amplitude distributions are shown for the lower TiO₂ surfaceinside the hole for a single unit cell that repeats in both the X and Ydirections. As before, the cross product of the E and H field componentsdescribes the instantaneous power density distribution responsible forfluorescent excitation at the lower surface. A 1 V/m electric field, 1A/m magnetic field plane wave at the resonant wavelength is used as theillumination source.

Similarly, the electromagnetic field distributions may be computed forthe upper TiO₂ surface of the unit cell. FIGS. 19A-19C plots the X, Yand Z components of the electric field amplitude in the XY planecorresponding to the upper surface of the structure of FIGS. 15A and 15Bat Z=433 nm for incident wavelength 632.5 nm. FIGS. 19D-19F plot of X, Yand Z components of the magnetic field amplitude in the same XY planerepresented in FIGS. 19A-19C for incident wavelength 632 nm. Note thatthe maximum amplitude of each field component, as indicated by the plotlegend, is substantially higher than those of the prior art designindicating that higher power density may be obtained at the surface ofthis device. In particular, note that a substantial Ez component hasappeared in contrast to the Ez amplitude of the 1D prior design.

FIG. 19G plots reflection efficiency as a function of wavelength modeledusing incident illumination polarized parallel to the X axis. Light withX axis polarization, incident on the design represented by FIG. 15,generates a resonance useful for label-free detection, with a width ofapproximately 12.5 nm and a maximum near 830 nm. The simulation can alsopredict the bulk refractive index shift coefficient, defined as delta(PWV)/delta(n), where delta (PWV) is the shift in the peak wavelengthvalue induced by a refractive index change of delta (n) in theenvironment above the sensor. This quantity indicates the sensitivity ofthe sensor to binding of a sample to the grating surface. The “Hole”design described by FIGS. 15A and 15B has a predicted bulk shiftcoefficient of 200, indicating that the structure will provide sensitivelabel-free performance.

B. Posts Embodiment Example

A 2-dimensional grating structure using a repeating unit cellcharacterized by a post will now be described with reference to FIGS.20-24.

FIGS. 20A and 20B are perspective and cross-sectional views,respectively, of a unit cell of 2-dimensional grating designcharacterized by periodic posts 220 formed in the sensor surface. Eachunit cell has one post 220. The posts 220 are raised projections in asubstrate material 102 (e.g., UV cured polymer) which is applied to abase sheet (not shown). A high index of refraction (e.g., TiO₂) coatingis applied to the projections and substrate as shown in the Figures. Thestructure is optimized for BIND (label-free) detection in a waterenvironment using light polarized in the X direction and optimized forER detection in an air mode, using light polarized in the Y direction.

The design of FIG. 20 was studied by RCWA computer simulation. While theprevious structure unit cell of FIG. 15 contained a “hole” regionsurrounded by regions at a higher plane in the z-direction, the gratingstructure of FIG. 20 contains a central “post” region, surrounded byregions at a lower plane in the z-direction. As before, the design ofFIG. 20 represents a BIND/ER combined structure that is optimized toprovide a BIND resonance in the near infrared (˜800-900 nm) wavelengthregion, while providing an ER at 632 nm for excitation of the Cy5fluorophore. In this example, the design again assumes a waterenvironment over the sensor during BIND measurement and an airenvironment over the sensor during ER measurement. These differingwavelength requirements for ER and BIND, engender selection of arectangular “post” unit cell. Thus, the unit cell may have differingdimensions in the X and Y directions. For example, the period in the Xdirection is 530 nm for the BIND wavelength, but is 414 nm in the Ydirection as required for the lower wavelength ER resonance. Thefabrication process again dictates that the high refractive indexdielectric thickness will be the same in the X and Y directions. Forfabrication simplicity, the design also has uniform grating depth. Thefabrication process will also result in rounding of the post corners,however the principal function of the design remains unchanged. Theexample of FIG. 20 is meant as an illustrative example not limiting inscope. The specific dimensions can of course vary.

FIGS. 21A and 21B graph the reflection efficiency as a function ofwavelength and incidence angle, respectively, for the structuresrepresented by FIGS. 20A and 20B when illuminated with light polarizedalong the Y axis. These figures, generated by RCWA, represent operationin an ER mode. FIG. 21A shows a resonance peak maximum reflection at 633nm and FIG. 21B shows a resonance across a range of incident angles whenilluminated at 633 nm.

The RCWA computer simulations may be used to determine the spatialdistribution of the amplitude of the electric field components (Ex, Ey,and Ez) and the magnetic field components (Hx, Hy, Hz) at the ERresonant wavelength. FIGS. 22A-22C plot the X, Y and Z components ofelectric field amplitude in the XY plane corresponding to the lowersurface of the structure of FIGS. 20A and 20B at Z=70 nm, for incidentwavelength 633 nm. FIGS. 22D-22F plot the X, Y and Z components of themagnetic field amplitude in the same XY plane represented in FIGS.22A-22C for incident wavelength 633 nm. As shown in FIG. 22, the fieldamplitude distributions are shown for the lower surfaces surrounding theposts for a single unit cell that repeats in the X and Y directions. Asbefore, the cross product of the E and H field components describes theinstantaneous power density distribution responsible for fluorescentexcitation at the lower surface. A 1 V/m electric field, 1 A/m magneticfield plane wave at the resonant wavelength again serves as theillumination source.

Similarly, the electromagnetic field distributions may be computed forthe upper TiO₂ surfaces of the unit cell. FIGS. 23A-23C plot the X, Yand Z components of the electric field amplitude corresponding to theupper surface of the structure of FIGS. 20A and 20B at Z=430 nm forincident wavelength 633 nm. FIGS. 23D-23F plot the X, Y and Z componentsof the magnetic field amplitude in the same XY plane represented byFIGS. 23A-C for incident wavelength 632 nm. Note that the maximummagnitude of the fields is again substantially higher than the prior artdesign (FIG. 10) for each of the field components, indicating thatpotentially higher power density may be obtained at the surface of thisdevice.

FIG. 24 plots reflection efficiency modeled using incident illuminationpolarized parallel to the X axis. Light with X-axis polarization,incident on the deign represented by FIG. 15, generates a resonanceuseful for label-free detection, with a width of approximately 8 nm andreflectance maximum near 805 nm. The simulations produce a shiftcoefficient (explained above) of 90. The embodiment of FIG. 20 is alsoexpected to provide sensitive label-free performance, though probablylower than the embodiment of FIG. 15. Comparison of amplitude values andshift coefficients between the two 2D designs suggests that furtheroptimization of the grating structure could emphasize the performance ofone detection mode as a tradeoff for reduced performance in the otherdetection mode.

The amount of amplification for ER detection relates to the powertransferred from the device structure to a distribution of fluorophoreson the sensor surface within at the excitation wavelength range of thefluorophore. The power density distribution of the sensor surface at theresonant wavelength, provided that the resonant wavelength falls withinthe excitation wavelength range, therefore provides a means forcomparing the sensitivity of different ER device designs. One can definethe cross product E (max)×H (max) as a field power or “magnificationfactor”. While a more thorough analysis of the intensity distribution ofthe evanescent field from the tops, bottoms, and sides of the structure,and a detailed integration of power density to account for differencesbetween higher and lower power regions would provide a more exactprediction of whether one device will function more effectively thananother, the product of the maximum amplitude of an E component with anorthogonal H component provides a very simple, rough way of comparingdesigns. Using RCWA analysis for the exposed upper and lower planes ofthe devices, the E×H magnification factor for the prior art design ofFIGS. 11A and 11B is 144. Conversely, for the “holes” unit cell designof FIGS. 15A and 15B, the E×H magnification factor is 6217, while forthe “posts” design of FIGS. 20A and 20B, the E×H magnification factor is5180. Based on this rough analysis, the ER aspects of the 2D gratingdesigns of FIGS. 15 and 20 appear to provide the potential for highersensitivity ER performance than the linear grating design of FIG. 10.Moreover, the designs of FIGS. 15 and 20 are expected, based on thecomputer simulations, to provide excellent sensitivity for label-freedetection, as explained above. Therefore, useful combined ER andlabel-free detection in a single device is achieved in theabove-described 2D grating embodiments.

C. Two-Level, 2-D Gratings

FIGS. 28A-C are three perspective views of yet another embodiment of aunit cell 500 for a biosensor grating structure constructed and designedfor a combined ER and label-free (BIND) detection. In order toappreciate some of the features of this structure, it will be useful torecapitulate on the design aspects pertinent to evanescent resonance(ER) and label-free (BIND) sensors. Such sensors differ in three basicdesign aspects, namely: resonance wavelength, resonance width, andgrating depth.

Resonance Wavelength

The ER sensor prefers resonance to occur in within a few (˜+/−2) nm ofthe excitation wavelength. Given that the excitation light generallycomes from a laser and has very narrow bandwidth, this requirementplaces high specificity on the wavelength location of the ER resonance.The BIND mode of operation does not have this limitation and may benefitfrom a resonance at another wavelength e.g. outside ambient lightingwavelength range or to separate the BIND signal spectrally from the ERexcitation source thereby eliminating potential overlapping detectionconflicts.

Resonance Width

The ER sensor must have a resonance wide enough for it to overlap theexcitation wavelength in the presence of variables such as biologicalcoating thickness and illumination numerical aperture. In practice, theER resonance should not have a full width at half maximum (FWHM) lessthan about 5 nm, and more preferably between 10 and 15 nm. On the otherhand, BIND sensitivity increases approximately as 1/sqrt (FWHM) becausepeak location uncertainty decreases as the peak width narrows.

Grating Depth

BIND sensors give greater resonance wavelength shift when morebiological material adheres to the grating. A deeper grating offers moresurface area for binding biological material. The ER effect does notnecessarily improve and may degrade as the ER grating depth increases.

The 2-D designs described previously have uniform grating depth (e.g. inthe post examples the height of the posts, or in the holes example thedepth of the holes). Selecting a single grating depth may involve acompromise between BIND and ER performance both in terms of peak widthand surface area, i.e. BIND PWV shift.

The design of the biosensor of FIG. 28A-C is a two-level,two-dimensional design. The specifics of the design will be discussedbelow in greater detail. This design maintains a narrow TM BINDresonance and high BIND shift performance, while simultaneous providinga wider TE ER resonance. Similar to previously described two-dimensionaldesigns, the BIND and ER gratings can have different periods and henceindependently determined resonance wavelengths.

This “two level” “comBIND” design of FIG. 28A-C comprises a multitude ofrepeating unit cells 500, each of which superimposes a relativelyshallow ER grating 502 extending in the X direction on a relatively deepBIND grating 504, extending in the Y direction. FIGS. 28A-29C depict one“unit cell” 500 for this design, which, when replicated in the XY planeforms the complete grating.

The unit cell 500 consists of a UV-cured polymer layer 524 which isapplied using a master grating wafer to a base substrate sheet such asPET film (not shown). The polymer layer 524 has the structure of theBIND grating 504, namely alternating low and high regions extending inthe Y direction. In the X direction, the grating also has alternatinglow and high regions, although the relative height of the high regioncompared to the low regions of the UV-cured polymer layer 524 in the Xdirection is much less than in the Y direction.

A TiO₂ (or alternatively SiO₂ or Ta₂O₅) layer 522 is deposited over theUV-cured polymer layer. This layer has uniform thickness in theillustrated embodiment. The layer 522 includes upper repeating surface506, 508, 510, and 512, and lower repeating surface 514, 516, 518 and519. The lower surfaces 514, 516, 518 and 519 are positioned over thetop surface of the UV-cured polymer layer. An air or water sample medium520 is placed in contact with the upper surfaces 506, 508, 510, 512 ofthe TiO₂ or SiO₂ layer 522.

As will be appreciated from inspection of FIGS. 28A-C, the “two-layer2-D” grating structure includes a relatively deep BIND grating 504 inthe Y dimension, characterized by upper and lower grating surfaces506/508 and 510/512, respectively. The BIND aspect of the unit cell thuspermits adding or more sample material and allows more material toadhere to the grating, permitting a greater resonance shift. The deepergrating in the BIND (Y direction) offers more surface area for bindingbiological material.

The ER grating 502 extending in the X direction, conversely, consists ofa relatively shallow grating pattern with high regions 506 and lowregions 508 (and also high region 510 and low region 512). In additionto providing good BIND detection capability, the grating is expected tosimultaneously provide a wider TE ER resonance with optimal width.

An apparent advantage of the design of FIGS. 28A-C is that the ER andBIND structures should operate independently. Hence, structuraldimensions optimized for either ER detection or BIND detection aloneshould work for the combination of the ER and BIND sensor of FIG. 28A-C.While the specific dimensions for a structure having the unit cell ofFIG. 28A-C is of course variable, in one representative embodiment theBIND grating 504 has a period of between about 260 and about 1500 nm,and the depth of the grating (distance between surfaces 506 and 510) isbetween 100 nm and about 3000 nm. For the ER grating 502, the period isbetween about 200 nm and about 1000 nm, and the depth (Z distancebetween surfaces 506 and 508, and 510 and 512) is between 10 nm andabout 300 nm.

The structure of FIG. 28A-C was simulated on a computer using RCWA andits simulated reflection spectrum obtained, both with and without theaddition of an ER grating structure in the X direction. FIG. 30 showsthe BIND grating spectrum without the ER grating 502 in the X direction(curve 592), and the combined ER and BIND grating spectrum (curve 590).Both spectra simulations include water on the surface of the biosensor.The addition of the ER grating over the BIND grating (curve 590) createsadditional resonance peaks (600 and 602) of width and locationappropriate for ER excitation of a CY5 fluorophore. Note that the ERgrating 502 (FIG. 28A-C) and curve 590 also enhances total surface areathus offering the potential for BIND shift improvement. BIND peakwavelength values are indicated at peaks 606 and 604 in FIG. 30.

Further Applications for ER+BIND Biosensor Using Inhibitors

Consider further the following inhibition binding scenario, in whichsome fraction of the substance to be detected, e.g., protein, bindsdirectly to the grating substrate (label-free) and some other fractionof the protein binds to an inhibitor having a fluorescent label. K_(s)and Ki are equilibrium binding constants for the substrate (K_(s)) andthe inhibitor (K_(i)).

Mathematical equations can be written defining values and relationshipsfor the concentrations and equilibrium binding constants (K) for thetypical inhibition binding scenario above:Ks=[protein˜substrate]/[protein]×[substrate] and  (1)K _(I)=[protein˜inhibitor]/[protein]×[inhibitor]  (2)Combining the two equations and rearranging terms, one can easily arriveat an equation for the fraction of protein bound by the substrate.

${{Fraction}\mspace{14mu}{of}\mspace{14mu}{protein}\mspace{14mu}{bound}\mspace{14mu}{by}\mspace{14mu}{the}\mspace{14mu}{substrate}} = \frac{1}{{2{PX}} - \left. \sqrt{}\left( {X^{2} - {4{PS}}} \right) \right.}$

-   Where X=(K_(s)+(K_(s)/K_(i))*I+S+P)-   K_(i) is the inhibition binding constant-   K_(s) is the substrate binding constant-   and P, S, and I are the concentrations of the protein, substrate,    and inhibitor, respectively.

In the setup of test reactions for new inhibitors, the operatorgenerally only measures the amount of substrate bound, using for examplesome fluorescent label, but is unable to directly quantify the inhibitorligand binding or the K_(I) at the same time (i.e., the ligand bindingis inferred). With two independent quantification methods provided bythe combined ER and label-free biosensor of this disclosure, and asimple test, all of the variables for the binding scenario describedabove can be known. In other words, a more complete understanding andcharacterization of the binding properties are obtained using a singletest using the inventive ER and label-free sensor.

The fluorescent label could be on the grating substrate surface or onthe inhibitor. This means that either the substrate or the inhibitorcould be label-less (label-free). A preferred embodiment uses a firstbinding molecule (that could the substrate of the biosensor) and asecond potential binding molecule, e.g., inhibitor molecule, that mayinfluence or compete with the binding of the biological substance (e.g.,protein) with the first binding molecule.

This technique of using inhibitors to influence binding reactions in alabel and label-free biosensor could be extended to encompass very tightbinding interactions whereby a known competitive inhibitor with a weakerbinding affinity could be employed to perturb/observe the much tighterbinding entity.

In general, examples of specific binding substances (samples) which maybe detected with the biosensor of this invention include nucleic acids,polypeptides, antigens, polyclonal antibodies, monoclonal antibodies,single chain antibodies (scFv), F(ab) fragments, F(ab′)2 fragments, Fvfragments, small organic molecules, cells, viruses, bacteria, polymers,peptide solutions, protein solutions, chemical compound librarysolutions, single-stranded DNA solutions, double stranded DNA solutions,combinations of single and double stranded DNA solutions, RNA solutionsand biological samples. Such biological samples could consists of, forexample, blood, plasma, serum, gastrointestinal secretions, homogenatesof tissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid,amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavagefluid, semen, lymphatic fluid, tears and prostatic fluid.

The biosensor described herein may be used to detect (a) binding ofcomponents any of these types of samples to the biosensor surface, (b)binding of the sample to another component of the sample, e.g., afluorophore in the sample, and (c) binding of the sample or samplecomponent to a second sample which is added to the sample. As an exampleof binding (b), the sensor surface may bind to some component of thesample, such as for example streptavidin-biotin or 6His, and thebiosensor may be used to detect the interaction of the bound componentof the sample with an additional grouping of components in the sample,such as a polymerase complex. In the latter example of binding (c), asample may have a component that is attached to the surface of thebiosensor and another component which specifically binds/attractsanother component(s) from a second sample that is placed on thebiosensor.

Embodiments Using Natural Fluorescence

As a further embodiment, the combined ER and label-free biosensor isparticularly useful for assays which utilize the natural fluorescence ofbiological molecules (i.e., without requiring the use of a boundfluorescence label), to make biophysical characterization measurementsof folding, stacking, and changes and rates of changes to these uponinteractions with other biological molecules and small test molecules.Such characterization measurements could be made with a boundfluorescence label, but such bound label is not necessarily required,especially for biological materials having an inherent fluorescenceproperty.

The following table sets forth fluorescence characteristics of proteinand nucleic acid constituents and coenzymes.

TABLE 1 Fluorescence Characteristics of Protein and Nucleic AcidConstituents and Coenzymes Absorption Fluorescence* Sensitivity ^(λ)max^(ε)max ^(λ)max ^(τ)F ^(ε)max ^(ø)F Substance Conditions (nm) ×10⁻³ (nm)^(ø)F (nsec) ×10⁻² Tryptophan H₂O, pH 7 280 5.6 348 0.20 2.6 11.Tyrosine H₂O, pH 7 274 1.4 303 0.14 3.6 2.0 Phenylalanine H₂O, pH 7 2570.2 282 0.04 6.4 0.08 Y base Yeast tRNA^(phe) 320 1.3 460 0.07 6.3 0.91Adenine H₂O, pH 7 260 13.4 321 2.6 × 10⁻⁴ <0.02 0.032 Guanine H₂O, pH 7275 8.1 329 3.0 × 10⁻⁴ <0.02 0.024 Cytosine H₂O, pH 7 267 6.1 313 0.8 ×10⁻⁴ <0.02 0.005 Uracil H₂O, pH 7 260 9.5 308 0.4 × 10⁻⁴ <0.02 0.004NADH H₂O, pH 7 340 6.2 470 0.019 0.40 1.2 *Values shown for ^(ø)F arethe largest usually observed. In a given case actual values can beconsiderably lower. Source: Charles R. Cantor and Paul R. Schimmel,parts 1-3 Biophysical Chemistry - The behavior and study of biologicalmolecules, W.H. Freeman and Company, New York, (1980), page 443.

This technique is especially important with nucleic acid polymers (DNA,RNA) (fluorescent nucleoside bases) stacking and hybridization, proteins(fluorescent amino acids phenylalanine, tryptophan, and tyrosine) andlipid membranes (enhancement and quenching effects upon incorporation offluorophores into their different compartmentalizations). In oneembodiment, the label-free BIND feature allows the quantification of theamount of sample material or ligand bound thereto and the ER featureallows the sensitive tracking of the biophysical change.

ER and BIND Biosensor with Additional Low Fluorescence SiO2 Layer toReduce Background Fluorescence

When the combined ER and BIND sensor of this disclosure is used fordetecting fluorescence from a bound fluorophore or natural fluorescence,it can be useful to reduce background fluorescence emitting from withinthe biosensor construction materials so as to be able to generatefluorescence measurements with a higher signal to noise ratio. One wayof accomplishing this is to deposit an additional layer of lowfluorescence SiO₂ material onto the UV-cured polymer layer and thendeposit the uppermost high index (e.g., TiO₂) layer onto the SiO₂ layer.Such a biosensor is shown in FIG. 30, and includes UV-cured polymerlayer 524 (which is bound to a substrate sheet, not shown), intermediateSiO₂ layer 700, and upper TiO₂ layer 522. A sample in either an air orwater-based medium is placed on the TiO₂ layer. The thickness of theadditional SiO₂ layer 700 will depend on such factors as the gratingduty cycle and required background fluorescence level, but generallywill be in the range of 500 to 5000 Angstroms.

The SiO₂ 700 layer preferably has low native fluorescence in response toincident radiation from light sources used to interrogate the biosensor.The fluorescence level in SiO₂ depends on the process by which it ismade, and its structure (amorphous vs. nanocrystalline), may also play arole. Full oxidation of the SiO₂ molecules in the layer (SiO₂), ascompared to say SiO_(1.95), also appears to be important in providing alayer with low fluorescence. Preferably, the SiO₂ layer is made by aprocess, and has a structure, which results in relatively low nativefluorescence.

One example of the use of SiO₂ intermediate layers would be in thestructure of FIG. 28A-C, wherein below the top TiO₂ layer 522, a layerof low fluorescence SiO₂ is applied over a UV-cured polymer layer. TheSiO₂ intermediate layer is of uniform thickness in both the X and Ydirections.

The additional layer of low fluorescence SiO₂ material may also bepresent in the other biosensors described previously in this document.

It will be appreciated that in FIGS. 28 and 30 and in other figuresshowing the upper high refractive index layer, TiO₂ is not a requiredmaterial for the upper layer 522. Ta₂O₅ (tantalum pentoxide) could workalso. TiO₂ has a higher refractive index than Ta₂O₅, which is why it isgenerally used for a high refractive index layer. When using Ta₂O₅, ittakes more physical thickness to achieve the same optical thickness. Inthe examples presented herein, the range of thickness for the upperlayer 522, whether using TiO₂ or Ta₂O₅, is between about 70 nm to 250nm.

In one further possible embodiment a hafnium oxide coating is applied toa biosensor as a high index of refraction layer, replacing TiO₂ orTa₂O₅. At infrared and visible wavelengths, TiO₂ or Ta₂O₅ have noabsorption. However, at low wavelengths, these materials all start toabsorb, and the absorption increases as one goes to lower and lowerwavelengths. This is a problem for resonant devices, because theabsorption has the effect of diminishing the resonance (i.e. no peakwill be measured, or the peak will be small). Hafnium oxide does nothappen to have absorption, even for wavelengths as low as 400 nm. Thethickness of the hafnium oxide coating and grating dimensions would beselected to yield resonance at the UV wavelengths of interest.

Use of ER Sensor with Time Resolved Fluorescence (TRF) and FluorescencePolarization (FP) Measurements

One further example of the novel uses of the present ER and BIND sensoris the use of the sensor for time resolved fluorescence (TRF) andfluorescence polarization measurements. TRF and FP polarization are twomethods that would benefit greatly from enhanced signal derived from thecombined label-free and ER device of this disclosure. These two methodsare especially useful for separating specific ER signals for bindingevents from background signals and from molecules not participating inthe binding event. The fluorophore would need to be matched to thewavelength enhancement capability of the sensor.

The invention of a method for using ER with FP and/or TRF offers a userthe opportunity to cleanly detect the presence of a fluorophore involvedin a binding event and discriminate such detection from background ornon-participating molecules, and with much higher sensitivity.

FP and TRF are widely used techniques in the diagnostic andpharmaceutical test protein/compound screening industries. See forexample U.S. Pat. Nos. 6,432,632, 6,207,397, 6,159,750, 6,448,018,6,455,861, 6,566,143, and 5,504,337, the contents of each of which areincorporated by reference herein. See also the patents of Budach et al.,U.S. Pat. Nos. 6,870,630 and 6,707,561, and the patents of Neuschäfer etal., U.S. Pat. Nos. 6,078,705 and 6,289,144. The methods are popularbecause they allow for discrimination of binding signals from unwantedsignals such as background and non-binding molecules. The currentmethods would have improved signal to noise ratios (sensitivity) andallow for reduced reagent consumption. These improvements will beespecially useful in the area of study known as proteomics wheresensitivity and reagent are limiting. In addition the typical sorts ofbiophysical determinations (folding, proximity to other molecules, size,etc.) would be enhanced as well.

By offering the potential of 50-100× improved sensitivity in determiningbinding interactions, these techniques also have several commercialadvantages, including reduced reagent consumption, and greaterconfidence limits in low signals

It is believed that any instrument that can measure FP from theunderside of a clear bottom plate would work for the ER/BIND biosensorsdescribed herein. There are instruments commercially available that canmeasure both FP and TRF, such as for example the Molecular DevicesSpectraMax M5 instrument. However, the instrument for detection of FPmeasurements does not have to be the same instrument that makes TRFmeasurements.

The physical characteristics of the ER biosensor that make it moreadvantageous for FP or TRF measurements, by providing a greatly improvedsignal and signal to noise ratio, is the grating structure of thebiosensor (as described at length in this document) that creates theresonance effect for enhanced signal. FP and TRF are recommended methodsfor pharmaceutical screening activities because they allow users toavoid unwanted signals coming from the natural fluorescence of thecompounds they are screening (i.e. noise/background much higher withthese compounds).

Spotting Process and Quality Control with ER and BIND Sensor

In one possible use of the biosensor of this invention, the ER and BINDsensor is used to analyze a sample in a multitude of locations, each ofsuch locations defining a microarray spot of about 10-500 microns indiameter.

The term “spot” refers to a small quantity of sample material which isplaced on the surface of the biosensor. One produces such spots bydepositing tiny droplets of target solution containing the sample on thesurface of the biosensor and letting them dry. The droplet sizedetermines spot size. Different processes exist for depositing thedroplet and such techniques are generally known in the art. In onepossible implementation, the sensor surface comprises many spots in anarray (i.e., a microarray of spots), most of which have differingcompositions.

One then applies a common test material of varying complexity to theentire array. In one example, this test material comprises afluorescently labeled test material. The sensor is then interrogatedwith light to look for binding between the test material and the spot.To analyze a biosensor in the form of a microarray with a multitude ofspots, one obtains an image of the biosensor showing all the spots (seethe imaging readout instrument in the embodiment of FIG. 25 discussedbelow) and uses image analysis techniques or separate capture ofspectrographic data from each spot to determine the signal (intensityand shift in peak wavelength value) from each spot. Each spot providesone data point. The image must have a resolution dimension (pixel size)below the spot size.

Process variability during spotting (e.g. DNA spotting) of samplematerials onto an assay surface leads to indeterminate results forsubsequent binding of test samples. Uncontrolled variation in thedensity of sample material, and among spots, significantly hindersquantitation of test material binding frequency by fluorescence signal.Quantifying the sample material by applying a second label is generallynot practical.

The combined ER and label-free biosensor of this disclosure overcomesthese problems. In particular, the BIND signal is used to quantify theamount of material in the spot rapidly in a non-destructive way, withoutthe use of fluorophores or other added labels. The BIND measurements aremade prior to exposure of the sample (spots) to the fluorophore-labeledtest material. The ER measurements are made after exposure of the sampleto the fluorophore-labeled test material. The BIND measurements thusprovide a normalizing quantity for the fluorescence signal from eachspot. A BIND image of the spotted array also provides informationregarding spot morphology and can thus serve a quality control functionfor the spotting process.

One representative embodiment will now be described in conjunction withFIGS. 31-33. FIG. 31 is an image of a microarray of spots deposited on agrating-based sensor (which may or may not be optimized for both ER andBIND measurements) at 7 micron resolution. The image is captured by aCCD camera (see the embodiment of FIG. 25 discussed below). The DNAspots may be imaged at various micron resolution (pixel size), such as 7micron, 15 micron and 30 micron. FIG. 32 is a graph of a peak shift forone of the spots of FIG. 31 due to presence of a DNA sample being placedon the sensor. The shift is indicated by movement of the peak wavelengthvalue to the right for the DNA spot. The amount of this shift is aquantitative measure of the quantity of DNA present on the spot. FIG. 33is an illustration of a row of spots and a corresponding graph of PWVshift (in nanometers) as a function of position along the row. Thevariation in PWV shift is a quantitative measure of the amount of DNAlocated on the spots in the row. The area on the sensor surface that ismissing the spot of DNA sample material is shown as the dark location inthe image of the row of spots), the missing spot also indicated by thelow region in the graph. The graph of FIG. 33 thus provides further afurther qualitative and quantitative measure of the amount of DNA on thespots in the array.

The spotting aspects of this disclosure thus provides a method foranalysis of thickness, uniformity and morphology of DNA spots (or ingeneral spots of any biological material such as RNA, protein,carbohydrates, peptides etc.) on nanostructured optical surfaces such asthose used in the BIND label-free technique. The method is suitable foruse both in nanostructured optical surfaces such as the combined ER andlabel free sensor described herein. It can also be used for analysis ofspots on a grating structure which is just designed for one technique orthe other.

Quality control of printed microarrays is important for bothmanufacturers and users of microarrays. The printed DNA spots often donot have a recognizable label or tag attached to them (such asfluorescent, quantum dot, radioactivity). This makes it difficult toimage printed spots for quality assurance quickly and reliably andquantitatively before an assay is performed. The variation in the amountof DNA printed on the spots has a profound effect on the outcome ofhybridization assays and is particularly critical in diagnostics relatedapplications. In view of these issues, the inventive spotting processaspect of this disclosure provides a non-destructive, non-contact methodto image as-printed DNA spots on nanostructured optical surfaces (sensorsurfaces). This invention can be used to ascertain the quality andreliability of DNA spot printing process and weed out defectivemicroarrays thereby reducing the cost of manufacturing such microarrays.This invention can also be used to normalize final results from labeledassays performed using these chips (via fluorescence etc) to amount ofmaterial originally spotted, providing information on bindingaffinity/efficiency not previously available by other means.

The methods of this disclosure are preferably performed in anon-contact, non destructive manner. That is, the spots are imaged viaoptical means and both quantitative and qualitative information isobtained as to the spots as explained in FIGS. 31-33. This method isbelieved superior over prior art methods which userandom-sequence-short-oligonucleotides that contain a labeled end (suchas a fluorescent tag) which weakly binds to the single stranded DNA onthe surface of the chip. The fluorescence signal from the bound spots isindicative of the amount of DNA present in the spots. These methods areoften plagued by problems such as dissociation at room temperature (dueto low melting point of random-labelled-oligo and DNA complex),streaking, spotting, non-specific binding to the substrate and therebyincreasing background and in some cases complete removal of the boundDNA from the spots owing to the use of detergents in these QC tests.Perkin Elmer has used reflective imaging to non-destructively determinethe presence and absence of spots using reflected laser light from saltcrystals present in the DNA spots. This is a purely qualitative (yes orno) method and does not work when the printing solution does not usesalt or if the salt crystals are washed off from the spotted array.

The functional advantages of the imaging technique of this aspect ofthis disclosure is that it provides a quantitative analysis of amount ofmaterial bound to structured optical surface. It is non-destructive andnon-contact based measurement method. Furthermore, the invention couldbe used by manufacturers and users of microarrays to ascertain thequality (uniformity, spot morphology and quantity of material bound) ofthe spots on the microarray surface. Further, the analysis takesapproximately one minute, instead of the considerably longer periodrequired in prior art. This allows analysis to be performed on everymicroarray, rather than smaller samples.

Readout Systems for Biosensors Combining Label-Free Detection andFluorescence Amplification (ER)

With the above description of combined ER and label-free biosensors inmind, this document will now describe several embodiments of a readoutand detection system useful for interrogating the sensor and acquiringboth label-free and ER data from a single binding site on the detector.

A first embodiment of a readout and detection system 300 is shownschematically in FIG. 25. The system 300 of FIG. 25 is an imagingreadout system. The biosensor 100 is designed to exhibit both a sharpresonant peak, in the optical spectrum, for label-free detection and ahigh electromagnetic field in the evanescent region of the biosensor forsignificant enhancement of fluorescence signal. The readout system readsout both of these effects, taking advantage of these biosensorproperties. This disclosure provides a novel imaging readout system withthe capability to measure either or both signals from the biosensor.

The biosensor 100, referred to herein as a “comBIND sensor” herein, isinterrogated optically from the bottom side of the sensor. On thetopside of the biosensor 100, the biosensor may be immersed in water oranother liquid, or it may be exposed to air. Any molecular or cellularbinding interaction, which the biosensor is designed to detect, takesplace on the topside of the biosensor 100. The biosensor 100 may be partof a larger assay device that includes liquid containing vessels, suchas for example a microwell plate having e.g., 8 columns of wells, eachrow containing 12 wells. The biosensor may also be a component of amicroarray slide. In the illustration of FIG. 25, a single well(detection site) 302 is shown in cross-section, it being understood thatdozens, hundreds or even thousands of such detection sites may bepresent.

The imaging readout and detection system 300 includes an ER light source340 in the form of a laser (e.g., HeNe laser), a broader spectrum BINDlight source 350 including as a halogen white light source or a LED 352,and a CCD camera system 338 serving as a common detector to capture bothER and label-free data in successive images. The system 300 includes anoptical beam combining subsystem that includes dichroic mirrors 364 and330 which serves to combine and direct incident light 372 from the lightsources 340 and 352 onto the biosensor. The dichroic mirror 330 collectssignal light for detection and directs it to a lens 336 where it isimaged by the CCD camera 338.

The light beam 370 present below the biosensor 100 consists ofillumination light 372 and reflected light 374. The reflected light 374includes direct reflection and fluorescent emission if there isfluorescent material present on the biosensor.

Signal detected by the CCD camera 338 through a lens system 336 isprocessed electronically or by computer algorithm to become BIND(label-free) data 380 or ER data 382. Such data may be stored,displayed, and analyzed on an analytical instrument such as a computeror workstation for the instrumentation shown in FIG. 25 (not shown, buthaving access to data 382 and 380) by the user of the readout system300. Furthermore, the combination of the BIND data 380 and the ER data382 allows the user to gain information on binding interactions or cellinteractions that is unique to the novel biosensor 100.

In the illustrated design, the optical components 340, 350 and 330 aredesigned to produce a single beam 372 of incident radiation and thebiosensor is moved in X and Y directions to thereby sequentially obtaindata from all the wells 302 or binding sites on the biosensor 100surface. Such motion may be produced by placing the biosensor 100 on anX-Y motion stage (not shown), of which persons skilled in the art arefamiliar. When a given well or binding site 302 is in position such thatthe well 302 is in registry with the beam 372, in one embodiment thelight sources 340 and 350 are operated in succession (or selectivelyallowed to direct radiation onto the biosensor) and first and secondimages are captured by the CCD camera 338, one an ER image and the othera BIND image. The successive collection of CCD images could befacilitated by use of the beam selection mechanism 360 (such as ashutter), which selectively allows light from either the source 340 orthe source 350 to pass to the dichroic mirror 330 and be reflected ontothe biosensor. Beam selection can also be done electronically, such asby electronically controlling the on and off times of the light sources340 and 350. Alternatively, both light sources could be activated at thesame time and the selection mechanism 360 operated to pass both beams sothat the incident beam 372 contains light from both sources. In thissituation, the CCD camera 338 would capture a single image containingboth ER and BIND information. Image processing techniques would then beapplied to the resultant image from the CCD camera 338 to extract theBIND and ER components of the composite image.

The ER light source 340 may be a laser, such as a helium-neon (HeNe)laser. The laser beam 341 further goes through a beam-conditioningdevice 342 such as a beam expander. The beam expander 342 expands asmall diameter laser beam into a large diameter laser beam. The outputbeam 343 is collimated and linearly polarized. The biosensor producesthe ER effect in response to incident light at a specific polarization.Polarization may be achieved by using a laser designed for producing alinearly polarized output laser beam.

The BIND (label-free) light source 350 may consist of a halogen or LEDlight source 352, and a monochromator 354 with a wavelength adjustmentmechanism 356. The light beam 353 emitted by the light source 352 isbroadband in nature, while the light beam 355 at the exit port of themonochromator 354 is monochromatic.

The output light beam 355 from the monochromator 354 is conditioned by abeam conditioning device 358, which may be a collimator. A mirror 365directs the light beam 349 from the output of the conditioning device358 to the dichroic mirror 364. The combined light from the lightsources 340 and 350 is shown at 366 where it is directed to the beamsplitting and combining assembly 330 which then directs it to the bottomsurface of the biosensor 100.

The BIND light source 350 may also consist of a tunable laser. In thatcase, the beam-conditioning device 358 is a beam expander. Note alsothat a tunable laser or flash lamp could serve as a single illuminationsource for both BIND and ER measurements.

In addition, since polarized light facilitates detection of a BINDsignal, there may be a polarizer within the light source 352 so that thelight 363 is linearly polarized. Alternatively, the light-directingelement 365 may be a polarizing beam splitter to transform a randomlypolarized light 359 into a linearly polarized light 363.

For detection of the laser excited fluorescence signal, the beamsplitting and combining assembly 330 incorporates a set of opticalfilters 332 and 334. Filter 332 is a dichroic filter that reflects thelaser light while transmitting fluoresced light from the sample. Filter332 also functions as a beamsplitter in the BIND wavelength range, whichis 830 nm to 900 nm in one preferred design. Filter 334 only allowstransmission of light within two wavelength ranges: laser excitedfluorescence and the BIND wavelength range. An imaging lens 336 may beused to collect the fluorescence light at the biosensor surface andfocus it on the focal plane of the CCD camera 338.

The design of FIG. 25 also includes rotation apparatus to rotate thebiosensor relative to the incident beam 372 for purposes of ERdetection. In one possible embodiment, a rotation device 331 is attachedto the beam splitting and combining assembly 330 and rotates theassembly 330 as indicated by the arrows (thereby providing for rotationof the incident beam about angle θ). In an alternative embodiment,rotation device 331 is omitted and instead a rotational device 333 isattached to the XY motion stage which operates to rotate the XY motionstage (and biosensor 100 mounted thereon) relative to the (fixed)incident beam 372, as indicated by the arrows to the left of device 333in FIG. 26.

Additional lenses, mirrors and optical filters may be incorporated intothe readout system to achieve desired performance. Properly designedoptical filters may be used to eliminate undesired cross-talk betweenBIND detection and ER detection. In addition, a beam selection mechanismin the form of electronic or mechanical shutters 360 may be used toproperly synchronize light illumination and detection of the twochannels, so that only one light source illuminates the biosensor at agiven time, to eliminate any cross-talk.

A significant advantage of the biosensor readout system described inFIG. 25 is that both BIND and ER data may be collectedly simultaneously(or in rapid succession) at the same biosensor location. High-resolutionimaging methods are useful for high content bioassays such as cell-basedassays or microarrays.

An integrating single point detector may replace the CCD camera 338. Inthat case, the system produces an image by synchronizing sensor motion,over the location of the incident radiation 372, with the detectoroutput.

Further details on use of a CCD camera to obtain ER data from abiosensor can be found in the technical literature, e.g., an article ofDieter Neuschäfer, Wolfgang Budach, et al., Biosensors & Bioelectronics,Vol. 18 (2003) p. 489-497, the contents of which are incorporated byreference herein.

A second embodiment of a readout and detection instrument 300 is shownin FIG. 26. Whereas the design of FIG. 25 is an imaging readout system,the design of FIG. 26 is not an imaging system. As before, the biosensor100 is designed to exhibit both a sharp resonant peak in opticalspectrum for label-free detection and a high electromagnetic field inthe evanescent region of the biosensor for significant enhancement offluorescence signal. Because of these properties of the biosensor, asystem that reads out both of these effects is required. FIG. 26describes an additional novel readout system that measures either orboth signals from the biosensor.

The biosensor 100 is interrogated optically from the bottom side at thelocation of a binding site (e.g., well 302). The topside of thebiosensor 100 may be immersed in water or another liquid, or may beexposed to air. Any biomolecular or cellular binding interaction, whichthe biosensor is designed to detect, takes place on the topside of thebiosensor. Any of the measurement systems described herein could alsoread the biosensor 100 from the top (binding side), if desired, usingappropriate focusing apparatus.

As noted above, the sensor 100 may be part of a larger assay device thatincludes liquid containing vessels, such as a microwell plate. Thebiosensor may also be a component of a microarray slide.

The readout system 300 includes a BIND light source 402, a BIND detector400, an ER light source 406, and an ER detector 404. An optical system430 serves to combine light from the sources 406 and 402 and direct suchlight as an incident beam 450 onto the bottom surface of the biosensor100. The optical system 430 further collects reflected light 452 fromthe biosensor and directs such reflected light to the detectors 400 and404. The optical system 430 consists of four light beam splitting andcombining components, one (432) for the BIND detector, one (434) for theBIND light source, one (436) for the ER detector, and one (438) for theER light source. The light beam 450 present below the biosensor consistsof incident light 452 and returned light 454. The returned light 454includes reflected light and fluorescent emission if there isfluorescent material present on the biosensor.

Signal detected by the BIND (label-free) detector 400 is processedelectronically or by computer algorithm to become BIND data 380, whichis stored, displayed, and analyzed on a computer (not shown) by the userof the readout system. Similarly, signal detected by the ER detector 404is processed and transformed into ER data 382 for the user. Furthermore,the combination of BIND data 380 and ER data 382 allows the user to gaininformation on binding interactions or cell interactions unique to thenovel biosensor.

The embodiment of FIG. 26 also includes either a rotation device 331 forrotating the incident light beam 452 relative to the biosensor 100, or arotation device 333 for rotating the XY stage and biosensor 100 relativeto the incident light 452. The placement of the rotation device 331 maybe such that the entire assembly 430 is rotated.

FIG. 27 depicts a readout system like that shown in FIG. 26, but withmore detail. The BIND detector 400 in this embodiment is a spectrometer.The BIND light source 402 may take the form of a tungsten halogen lightbulb or a light emitting diode (LED). The ER detector 404 comprises aphoto-detector, such as a photomultiplier tube (PMT). The ER lightsource 406 is preferably a laser producing a beam with a wavelengthwithin the excitation band of the fluorophore for use with thebiosensor, such as a helium-neon (HeNe) laser for excitation of the CY5fluorophore.

The light beam 456, emitted from the BIND light source 42, is nominallycollimated light for detection of the BIND signal. Thus, the lightsource 402 may incorporate a collimation lens. In addition, since use ofpolarized light improves detection of the BIND signal, the light source402 may also incorporate a polarizer so that the light 456 is polarized.Alternatively, the beamsplitter 434 may be a polarizing beamsplitter totransform a randomly polarized light 456 into a linearly polarized lightincident on the bottom of the biosensor 100.

The light beam 458 from the laser 406 is a collimated beam. ERperformance improves if the laser beam has linear polarization.Polarization may be achieved by using a laser designed to produce alinearly polarized beam.

For detection of the laser-excited fluorescence signal from thebiosensor 100, the beam splitting and combining assembly 436 includes aset of optical filters 472, 474, and 476. Filter 472 allows the incidentlaser beam 458 from the laser source 406 to transmit through the beamsplitting and combining element 436. Filter 476 is a dichroic filterthat transmits the laser light from the source 406 while reflecting thefluoresced light from the biosensor 100 in the direction of the detector404. Filter 474 only allows transmission of fluoresced light to bedirected at the photo-detector 404. An imaging lens 478 may be used tocollect the fluorescence from the biosensor surface more efficiently.

Additional lenses, mirrors and optical filters may be incorporated intothe readout system to achieve desired performance. Properly designedoptical filters may be used to eliminate undesired cross-talk betweenBIND detection and ER detection. In addition, electronic or mechanicalshutters may be used to properly synchronize light illumination anddetection of the two channels, so that only one light source illuminatesthe biosensor at a given time, to eliminate any cross-talk.

As with the case with the design of FIG. 25, the optical components ofthe design of FIGS. 26 and 27 can be constructed and arranged to producea beam 452 of incident radiation at one location while an XY motionstage moves the biosensor in X and Y directions to thereby sequentiallyobtain data from all the wells 302 or binding sites on the biosensor 100surface. In one embodiment the light sources 402 and 406 operate insuccession generating data successively at the detectors 400 and 404from the given well or binding site 302 currently illuminated by beam452. The successive collection of ER and BIND data could be facilitatedby use of beam selection mechanism (such as a shutter), not shown inFIG. 26, or by electronic control of the light sources 402 and 406.Alternatively, both light sources could be activated at the same time sothat the incident beam 452 contains light from both sources. In thissituation, the detectors 400 and 404 obtain data simultaneously.

A significant advantage of the biosensor readout system described hereis that both BIND and ER data may be collectedly simultaneously (or inrapid succession) at the same biosensor location. The BIND detector 400and the ER detector 404 may be integrating detectors that obtain datafrom a single point per binding site or well but over a relatively largearea, or imaging detectors, such as CCD detectors, that collect datapixel by pixel at a user specified resolution. High-resolution imagingmethods are useful for high content bioassays such as cell-based assaysor microarrays.

Furthermore, the optical structure of FIGS. 25, 26 and 27 can bereplicated such that multiple wells or binding sites on the biosensor100 can be interrogated and detected at the same time, e.g., takingadvantage of the concepts shown in FIG. 3.

Readout System with Single Light Generating Source

It will be noted that the embodiments of FIG. 25 includes separate lightsources 350 and 340 for BIND and ER measurements, respectively, and thatthe embodiment of FIG. 26 includes separate light sources 402 and 406for BIND and ER measurements, respectively. In one possible variation, asingle light source may be used for both BIND and ER measurements. Thislight source could take several possible forms, such as the form of atunable laser, or a broad spectrum high intensity flash lamp. The outputfrom the light source is optionally collimated, expanded with a beamexpander (if a tunable laser is used as the source), passed through amonochromator or filter stage (if a flash lamp is used), and thendirected to the surface of the biosensor. The optics used for detectionof ER and BIND signals can take the form of the apparatus shown in FIGS.25-27 and described above.

In one embodiment, to obtain both BIND and ER data from a given bindingsite in the biosensor, the light source can be activated twice in orderto select different wavelength ranges for illumination (one for BIND andone for ER). For example, the BIND detector obtains BIND data in thefirst activation of the light source and the ER detector obtains ER datain the second activation.

In one possible variation, in the case of a broad band source (i.e.Xenon flash lamp), one can also illuminate the biosensor with a broadspectrum and simultaneously collect BIND and ER data by splitting thereturn signal and diverting it through two different filter stages. Forsimultaneous BIND/ER illumination/collection, one needs to illuminatethe plate at some angle of incidence. The specular (direct) reflectioncomponent contains a BIND peak, and its spectral position is determinedby a monochrometer. A lens system, or even an integrating sphere,collecting only light leaving the surface at angles other than theincidence angle, will provide a relatively clean ER signal after passingthrough a filter or monochrometer that selects the emission range forthe fluorophore of interest.

While a number of exemplary aspects and embodiments have been discussedabove, those of skill in the art will recognize certain modifications,permutations, additions and sub-combinations thereof as being present inthe disclosure. It is therefore intended that the following claims andclaims hereafter introduced are interpreted to include all suchmodifications, permutations, additions and sub-combinations as arewithin their true spirit and scope.

In the claims, the term “evanescent resonance (ER) detection” or“evanescent resonance (ER) detection mode” is intended to encompass thedetection of fluorescence, phosphorescence, chemi-luminescence,electroluminescence, or other type of luminescence, for example asdescribed in Budach et al., U.S. Pat. No. 6,707,561. Such luminescencecould be attributable to native luminescence of the sample material orto a bound substance, e.g., fluorescence label, or quantum dots(luminescent metals). Such bound substance may be bound to the samplebeing tested, the surface of the biosensor, or both.

1. A sensor for testing a sample placed on the sensor, comprising: asubstrate having a periodic surface grating structure wherein theperiodic grating structure is constructed in a manner designed forboth 1) optical interrogation of the sensor with light in an evanescentresonance (ER) detection mode, and 2) optical interrogation of thesensor with light in a label-free detection mode.
 2. The sensor of claim1, wherein the sample is in an air medium, and further comprising adetection instrument which includes at least one light source foroptical interrogation of the sensor in the ER and label-free detectionmodes, and wherein the light from the at least one light source has apolarization perpendicular to the grating structure.
 3. The sensor ofclaim 1, wherein the sample is in a liquid medium, and furthercomprising a detection instrument which includes at least one lightsource for optical interrogation of the sensor and wherein the lightfrom the at least one light source has a polarization parallel to thegrating structure.
 4. The sensor of claim 2, wherein the light from theat least one light source has a wavelength selected to activate aluminescent material bound to (a) the sample, (b) the surface of thesensor, or (c) both (a) and (b).
 5. The sensor of claim 2, wherein thelight from the at least one light source interrogates the sensor at ornear normal incidence.
 6. The sensor of claim 2, wherein the detectioninstrument includes a software module calculating a peak wavelengthvalue of reflected light for the sample in a label-free detection mode.7. The sensor of claim 3, wherein the light from the at least one lightsource has a wavelength selected to activate a luminescent materialbound to (a) the sample, (b) the surface of the sensor, or (c) both (a)and (b).
 8. The sensor of claim 3, wherein the light from the at leastone light source interrogates the sensor at or near normal incidence. 9.The sensor of claim 3, wherein the detection instrument includes asoftware module calculating a peak wavelength value of reflected lightfor the sample in a label-free detection mode.
 10. The sensor of claim1, wherein the grating structure comprises a two-dimensional periodicgrating structure, the periodic grating structure is periodic in firstand second dimensions, and wherein the first and second dimensions aremutually orthogonal.
 11. The sensor of claim 10, wherein the firstdimension of the periodic grating structure comprises a gratingstructure designed for label-free detection, the second dimension of theperiodic grating structure comprises a grating structure designed for ERdetection, and wherein the depth of the grating structure in the firstdimension is substantially greater than the depth of the gratingstructure in the second dimension.
 12. The sensor of claim 11, whereinthe grating structure further comprises a laminate comprising asubstrate, a layer having a grating structure bonded to the substrate,an SiO₂ layer deposited on the layer having the grating structure, and alayer of relatively high index of refraction material deposited on theSiO₂ layer.
 13. The sensor of claim 11, wherein the grating structure inthe first dimension has a period of between 260 and about 1500 nm and adepth of the grating between about 100 nm and about 3000 nm, and whereinthe grating structure in the second dimension has a period of betweenabout 200 nm and about 1000 nm, and wherein the depth of the grating inthe second dimension is between about 10 nm and about 300 nm.
 14. Thesensor of claim 12, wherein the SiO₂ layer comprises a layer having adepth of between about 500 and about 5000 Angstroms.
 15. The sensor ofclaim 12, further comprising a luminescent material bound to the layerof high index of refraction material.
 16. The sensor of claim 11,wherein the grating structure comprises an array of unit cells having atwo-level, two-dimensional form.
 17. The sensor of claim 1, wherein thegrating structure has a grating depth to half period ratio of between0.6 and 1.2.
 18. The sensor of claim 1, wherein the grating structurecomprise array of unit cells each having a two-level, two-dimensionalform comprising a first periodic grating structure of alternating highand low regions extending in a first dimension constructed in a mannerdesigned for label-free detection and having a second periodic gratingstructure having alternating high and low regions superimposed on thefirst periodic grating structure extending in a second dimensionorthogonal to the first dimension constructed in a manner designed forER detection mode.
 19. The sensor of claim 10, wherein the gratingstructure comprises an array of two dimensional unit cells eachcomprising a hole in the surface of the biosensor.
 20. The sensor ofclaim 10, wherein the grating structure comprises an array of unit cellseach comprising a post projecting from the surface of the biosensor. 21.The sensor of claim 10, wherein the grating structure comprises acheckerboard configuration with alternating high and low regions. 22.The sensor of claim 10, wherein the grating structure in the first andsecond orthogonal dimensions has a different periodicity in the firstand second dimensions, the periodicity in the first dimension isconstructed so as to provide a broad resonance for incident light fromthe first light source at or near normal incidence with a wavelengthtuned to excite a fluorophore associated with the sample, and theperiodicity in the second dimension is constructed so as to yield asharp resonance for illumination with light from the second light sourcein the near infra-red portion of the spectrum.
 23. The sensor of claim1, further comprising at least one sample placed on the biosensor,wherein the sample is selected from the group of samples consisting ofmolecules having a molecular weight of less than 1000 daltons, moleculeswith a molecular weight of between 1000 and 10,000 daltons, amino acids,proteins, nucleic acids, lipids, carbohydrates, nucleic acid polymers,viral particles, viral components, cellular components, and extracts ofviral or cellular components, polypeptides, antigens, polyclonalantibodies, monoclonal antibodies, single chain antibodies (scFv), F(ab)fragments, F(ab′)2 fragments, Fv fragments, small organic molecules,cells, viruses, bacteria, polymers, peptide solutions, proteinsolutions, chemical compound library solutions, single-stranded DNAsolutions, double stranded DNA solutions, combinations of single anddouble stranded DNA solutions, RNA solutions and biological samples. 24.The sensor of claim 23, wherein the biological samples comprise samplesselected from the group of consisting of blood, plasma, serum,gastrointestinal secretions, homogenates of tissues or tumors, synovialfluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinalfluid, peritoneal fluid, lung lavage fluid, semen, lymphatic fluid,tears and prostatic fluid.
 25. The sensor of claim 1, further comprisinga sample placed on the sensor, wherein the ER detection mode detectsnatural fluorescence of the sample.
 26. The sensor of claim 1, furthercomprising a sample placed on the biosensor, wherein a fraction of thesample is bound to an inhibitor.
 27. The sensor of claim 26, wherein theinhibitor is bound to a fluorescent label.
 28. The sensor of claim 26,wherein the sample comprises a protein.